Radioactive emission detector equipped with a position tracking system

ABSTRACT

A radioactive emission probe in communication with a position tracking system and the use thereof in a variety of systems and methods of medical imaging and procedures, are provided. Specifically, wide-aperture collimation-deconvolution algorithms are provided, for obtaining a high-efficiency, high resolution image of a radioactivity emitting source, by scanning the radioactivity emitting source with a probe of a wide-aperture collimator, and at the same time, monitoring the position of the radioactive emission probe, at very fine time intervals, to obtain the equivalence of fine-aperture collimation. The blurring effect of the wide aperture is then corrected mathematically. Furthermore, an imaging method by depth calculations is provided, based on the attenuation of photons of different energies, which are emitted from the same source, coupled with position monitoring.

RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.10/616,307 filed on Jul. 10, 2003, which claims the benefit of priorityunder 35 USC §119(e) of U.S. Provisional Patent Application No.60/394,936 filed on Jul. 11, 2002.

U.S. patent application Ser. No. 10/616,307 is also acontinuation-in-part (CIP) of U.S. patent application Ser. No.10/343,792 filed on Feb. 4, 2003, which is a National Phase of PCTPatent Application No. PCT/IL01/00638 filed on Jul. 11, 2001, which is acontinuation-in-part (CIP) of U.S. patent application Ser. No.09/727,464 filed on Dec. 4, 2000, now U.S. Pat. No. 7,826,889, which isa continuation-in-part (CIP) of U.S. patent application Ser. No.09/714,164 filed Nov. 17, 2000, now abandoned, which is acontinuation-in-part (CIP) of U.S. patent application Ser. No.09/641,973 filed Aug. 21, 2000, now U.S. Pat. No. 8,489,176.

PCT Patent Application No. PCT/IL01/00638 also claims the benefit ofpriority under 35 USC §119(e) of U.S. Provisional Patent Application No.60/286,044 filed on Apr. 25, 2001.

The contents of the above applications are all incorporated by referenceas if fully set forth herein in their entirety.

FIELD AND BACKGROUND OF THE INVENTION

The present invention relates to a radioactive emission probe equippedwith a position tracking system. More particularly, the presentinvention relates to the functional integration of a radioactiveemission probe equipped with a position tracking system as above withmedical imaging modalities and (or) with guided minimally-invasivesurgical instruments. The present invention is therefore useful forcalculating the position of a concentrated radiopharmaceutical in thebody in positional context of imaged portions of the body, whichinformation can be used, for example, for performing an efficientminimally invasive surgical procedure. The present invention furtherrelates to a surgical instrument equipped with a position trackingsystem and a radioactive emission probe for fine in situ localizationduring resection and (or) biopsy procedures, which surgical instrumentis operated in concert with other aspects of the invention.

The use of minimally invasive surgical techniques has dramaticallyaffected the methods and outcomes of surgical procedures. Physicallycutting through tissue and organs to visually expose surgical sites inconventional “open surgical” procedures causes tremendous blunt traumaand blood loss. Exposure of internal tissues and organs in this manneralso dramatically increases the risk of infection. Trauma, blood loss,and infection all combine to extend recovery times, increase the extentof complications, and require a more intensive care and monitoringregimen. The result of such open surgical procedures is more pain andsuffering, higher procedural costs, and greater risk of adverseoutcomes.

In sharp contrast, minimally invasive surgical procedures cause littleblunt trauma or blood loss and minimize the risk of infection bymaintaining the body's natural barriers to infection substantiallyintact. Minimally invasive surgical procedures result in faster recoveryand cause fewer complications than conventional, open, surgicalprocedures. Minimally invasive surgical procedures, such aslaparoscopic, endoscopic, or cystoscopic surgeries, have replaced moreinvasive surgical procedures in all areas of surgical medicine. Due totechnological advancements in areas such as fiber optics, micro-toolfabrication, imaging and material science, the physician performing theoperation has easier-to-operate and more cost-effective tools for use inminimally invasive procedures. However, there still exist a host oftechnical hurdles that limit the efficacy and increase the difficulty ofminimally invasive procedures, some of which were overcome by thedevelopment of sophisticated imaging techniques. As is further detailedbelow, the present invention offers further advantages in this respect.

Radionuclide imaging is one of the most important applications ofradioactivity in medicine. The purpose of radionuclide imaging is toobtain a distribution image of a radioactively labeled substance, e.g.,a radiopharmaceutical, within the body following administration thereofto a patient. Examples of radiopharmaceuticals include monoclonalantibodies or other agents, e.g., fibrinogen or fluorodeoxyglucose,tagged with a radioactive isotope, e.g., ^(99M)technetium, ⁶⁷gallium,²⁰¹thallium, ¹¹¹indium, ¹²³iodine, ¹²⁵iodine and ¹⁸fluorine, which maybe administered orally or intravenously. The radiopharmaceuticals aredesigned to concentrate in the area of a tumor, and the uptake of suchradiopharmaceuticals in the active part of a tumor, or other pathologiessuch as an inflammation, is higher and more rapid than in the tissuethat neighbors the tumor. Thereafter, a radiation emission detector,typically an invasive detector or a gamma camera (see below), isemployed for locating the position of the active area. Anotherapplication is the detection of blood clots with radiopharmaceuticalssuch as ACUTECT from Nycomed Amersham for the detection of newly formedthrombosis in veins, or clots in arteries of the heart or brain, in anemergency or operating room. Yet other applications include radioimagingof myocardial infarct using agents such as radioactive anti-myosinantibodies, radioimaging specific cell types using radioactively taggedmolecules (also known as molecular imaging), etc.

The distribution image of the radiopharmaceutical in and around a tumor,or another body structure, is obtained by recording the radioactiveemission of the radiopharmaceutical with an external radiation detectorplaced at different locations outside the patient. The usual preferredemission for such applications is that of gamma rays, which emission isin the energy range of approximately 20-511 KeV. When the probe isplaced in contact with the tissue, beta radiation and positrons may alsobe detected.

The first attempts at radionuclide “imaging” were in the late 1940's. Anarray of radiation detectors was positioned mechanically on a matrix ofmeasuring points around the head of a patient. Alternatively, a singledetector was positioned mechanically for separate measurements at eachpoint on the matrix.

A significant advance occurred in the early 1950's with the introductionof the rectilinear scanner by Ben Cassen. With this instrument, thedetector was scanned mechanically in a predetermined pattern over thearea of interest.

The first gamma camera capable of recording all points of the image atone time was described by Hal Anger in 1953. Anger used a detectorcomprising a NaI(Tl) screen and a sheet of X-ray film. In the late1950's, Anger replaced the film screen with a photomultiplier tubeassembly. The Anger camera is described in Hal O. Anger, “Radioisotopecamera in Hine GJ”, Instrumentation in Nuclear Medicine, New York,Academic Press 1967, chapter 19. U.S. Pat. No. 2,776,377 to Anger,issued in 1957, also describes such a radiation detector assembly.

U.S. Pat. No. 4,959,547 to Carroll et al. describes a probe used to mapor provide imaging of radiation within a patient. The probe comprises aradiation detector and an adjustment mechanism for adjusting the solidangle through which radiation may pass to the detector, the solid anglebeing continuously variable. The probe is constructed so that the onlyradiation reaching the detector is that which is within the solid angle.By adjusting the solid angle from a maximum to a minimum while movingthe probe adjacent the source of radiation and sensing the detectedradiation, one is able to locate the probe at the source of radiation.The probe can be used to determine the location of the radioactivity andto provide a point-by-point image of the radiation source or data formapping the same.

U.S. Pat. No. 5,246,005 to Carroll et al. describes a radiation detectoror probe, which uses statistically valid signals to detect radiationsignals from tissue. The output of a radiation detector is a series ofpulses, which are counted for a predetermined amount of time. At leasttwo count ranges are defined by circuitry in the apparatus and the countrange which includes the input count is determined. For each countrange, an audible signal is produced which is audibly discriminable fromthe audible signal produced for every other count range. The mean valuesof each count range are chosen to be statistically different, e.g., 1,2, or 3 standard deviations, from the mean of adjacent lower or highercount ranges. The parameters of the audible signal, such as frequency,voice, repetition rate, and (or) intensity are changed for each countrange to provide a signal which is discriminable from the signals of anyother count range.

U.S. Pat. No. 5,475,219 to Olson describes a system for detecting photonemissions wherein a detector serves to derive electrical parametersignals having amplitudes corresponding with the detected energy of thephoton emissions and other signal generating events. Two comparatornetworks employed within an energy window, which define a function todevelop an output, L, when an event-based signal amplitude is equal toor above a threshold value, and to develop an output, H, when suchsignal amplitude additionally extends above an upper limit. Improvedreliability and accuracy is achieved with a discriminator circuit which,in response to these outputs L and H, derives an event output upon theoccurrence of an output L in the absence of an output H. Thisdiscriminator circuit is an asynchronous, sequential, fundamental modediscriminator circuit with three stable states.

U.S. Pat. Nos. 5,694,933 and 6,135,955 to Madden et al. describe asystem and method for diagnostic testing of a structure within apatient's body that has been provided with a radioactive imaging agent,e.g., a radiotracer, to cause the structure to produce gamma rays,associated characteristic x rays, and a continuum of Compton-scatteredphotons. The system includes a radiation receiving device, e.g., ahand-held probe or camera, an associated signal processor, and ananalyzer. The radiation receiving device is arranged to be locatedadjacent the body and the structure for receiving gamma rays andcharacteristic X-rays emitted from the structure and for providing aprocessed electrical signal representative thereof. The processedelectrical signal includes a first portion representing thecharacteristic X-rays received and a second portion representing thegamma rays received. The signal processor removes the signalcorresponding to the Compton-scattered photons from the electricalsignal in the region of the full-energy gamma ray and the characteristicX-ray. The analyzer is arranged to selectively use the X-ray portion ofthe processed signal to provide near-field information about thestructure, to selectively use both the X-ray and the gamma-ray portionsof the processed signal to provide near-field and far-field informationabout the structure, and to selectively use the gamma-ray portion of theprocessed signal to provide extended field information about thestructure.

U.S. Pat. No. 5,732,704 to Thurston et al. describes a method foridentifying a sentinel lymph node located within a grouping of regionalnodes at a lymph drainage basin associated with neoplastic tissuewherein a radiopharmaceutical is injected at the situs of the neoplastictissue. This radiopharmaceutical migrates along a lymph duct towards thedrainage basin containing the sentinel node. A hand-held probe with aforwardly disposed radiation detector crystal is maneuvered along theduct while the clinician observes a graphical readout of count rateamplitudes to determine when the probe is aligned with the duct. Theregion containing the sentinel node is identified when the count rate atthe probe substantially increases. Following surgical incision, theprobe is maneuvered utilizing a sound output in connection withactuation of the probe to establish increasing count rate thresholdsfollowed by incremental movements until the threshold is not reached andno sound cue is given to the surgeon. At this point of the maneuveringof the probe, the probe detector will be in adjacency with the sentinelnode, which then may be removed.

U.S. Pat. No. 5,857,463 to Thurston et al. describes further apparatusfor tracking a radiopharmaceutical present within the lymph duct and forlocating the sentinel node within which the radiopharmaceutical hasconcentrated. A smaller, straight, hand-held probe is employed carryingtwo hand actuable switches. For tracking procedures, the probe is movedin an undulatory manner, wherein the location of theradiopharmaceutical-containing duct is determined by observing a graphicreadout. When the region of the sentinel node is approached, a switch onthe probe device is actuated by the surgeon to carry out a sequence ofsquelching operations until a small node locating region is defined.

U.S. Pat. No. 5,916,167 to Kramer et al. and U.S. Pat. No. 5,987,350 toThurston describe surgical probes wherein a heat-sterilizable andreusable detector component is combined with a disposable handle andcable assembly. The reusable detector component incorporates a detectorcrystal and associated mountings along with preamplifier components.

U.S. Pat. No. 5,928,150 to Call describes a system for detectingemissions from a radiopharmaceutical injected within a lymph ductwherein a hand-held probe is utilized. When employed to locate sentinellymph nodes, supplementary features are provided including a functionfor treating validated photon event pulses to determine count rate levelsignals. The system includes a function for count-rate based ranging aswell as an adjustable threshold feature. A post-threshold amplificationcircuit develops full-scale aural and visual outputs.

U.S. Pat. Nos. 5,932,879 and 6,076,009 to Raylman et al. describe anintraoperative system for preferentially detecting beta radiation overgamma radiation emitted from a radiopharmaceutical. The system hasion-implanted silicon charged-particle detectors for generating signalsin response to received beta particles. A preamplifier is located inproximity to the detector filters and amplifies the signal. The probe iscoupled to a processing unit for amplifying and filtering the signal.

U.S. Pat. No. 6,144,876 to Bouton describes a system for detecting andlocating sources of radiation, with particular applicability tointeroperative lymphatic mapping (ILM) procedures. The scanning probeemployed with the system performs with both an audible as well as avisual perceptive output. A desirable stability is achieved in thereadouts from the system through a signal processing approach whichestablishes a floating or dynamic window analysis of validated photonevent counts. This floating window is defined between an upper edge anda lower edge. The values of these window edges vary during the analysisin response to compiled count sum values. In general, the upper andlower edges are spaced apart a value corresponding with about fourstandard deviations.

To compute these count sums, counts are collected over successive shortscan intervals of 50 milliseconds and the count segments resultingtherefrom are located in a succession of bins within a circular buffermemory. The count sum is generated as the sum of the memory segmentcount values of a certain number of the bins or segments of memory.Alteration of the floating window occurs when the count sum eitherexceeds its upper edge or falls below its lower edge. A reported mean,computed with respect to the window edge that is crossed, is developedfor each scan interval which, in turn, is utilized to derive a meancount rate signal. The resulting perceptive output exhibits a desirablestability, particularly under conditions wherein the probe detector isin a direct confrontational geometry with a radiation source.

U.S. Pat. No. 5,846,513 teaches a system for detecting and destroyingliving tumor tissue within the body of a living being. The system isarranged to be used with a tumor localizing radiopharmaceutical. Thesystem includes a percutaneously insertable radiation detecting probe,an associated analyzer, and a percutaneously insertable tumor removinginstrument, e.g., a resectoscope. The radiation detecting probe includesa needle unit having a radiation sensor component therein and a handleto which the needle unit is releasably mounted. The needle is arrangedto be inserted through a small percutaneous portal into the patient'sbody and is movable to various positions within the suspected tumor todetect the presence of radiation indicative of cancerous tissue. Theprobe can then be removed and the tumor removing instrument insertedthrough the portal to destroy and (or) remove the cancerous tissue. Theinstrument not only destroys the tagged tissue, but also removes it fromthe body of the being so that it can be assayed for radiation to confirmthat the removed tissue is cancerous and not healthy tissue. Acollimator may be used with the probe to establish the probe's field ofview.

The main limitation of the system is that once the body is penetrated,scanning capabilities are limited to a translational movement along theline of penetration.

An effective collimator for gamma radiation must be several mm inthickness and therefore an effective collimator for high energy gammaradiation cannot be engaged with a fine surgical instrument such as asurgical needle. On the other hand, beta radiation is absorbed mainlydue to its chemical reactivity after passage of about 0.2-3 mm throughbiological tissue. Thus, the system described in U.S. Pat. No. 5,846,513cannot efficiently employ high energy gamma detection becausedirectionality will to a great extent be lost and it also cannotefficiently employ beta radiation because too high proximity to theradioactive source is required, whereas body tissue limits the degree ofmaneuvering the instrument.

The manipulation of soft tissue organs requires visualization (imaging)techniques such as computerized tomography (CT), fluoroscopy (X-rayfluoroscopy), magnetic resonance imaging (MRI), optical endoscopy,mammography or ultrasound which distinguish the borders and shapes ofsoft tissue organs or masses. Over the years, medical imaging has becomea vital part in the early detection, diagnosis and treatment of cancerand other diseases. In some cases medical imaging is the first step inpreventing the spread of cancer through early detection and in manycases medical imaging makes it possible to cure or eliminate the canceraltogether via subsequent treatment.

An evaluation of the presence or absence of tumor metastasis or invasionhas been a major determinant for the achievement of an effectivetreatment for cancer patients. Studies have determined that about 30% ofpatients with essentially newly diagnosed tumor will exhibit clinicallydetectable metastasis. Of the remaining 70% of such patients who aredeemed “clinically free” of metastasis, about one-half are curable bylocal tumor therapy alone. However, some of these metastasis or evenearly stage primary tumors do not show with the imaging tools describedabove. Moreover often enough the most important part of a tumor to beremoved for biopsy or surgically removed is the active, i.e., growingpart, whereas using only conventional imaging cannot distinguish thisspecific part of a tumor from other parts thereof and (or) adjacent nonaffected tissue.

A common practice in order to locate this active part is to mark it withradioactivity tagged materials generally known as radiopharmaceuticals,which are administered orally or intravenously and which tend toconcentrate in such areas, as the uptake of such radiopharmaceuticals inthe active part of a tumor is higher and more rapid than in theneighboring tumor tissue. Thereafter, a radiation emission detector,typically an invasive detector, is employed for locating the position ofthe active area.

Medical imaging is often used to build computer models which allowdoctors to, for example, guide exact radiation in the treatment ofcancer, and to design minimally-invasive or open surgical procedures.Moreover, imaging modalities are also used to guide surgeons to thetarget area inside the patient's body, in the operation room during thesurgical procedure. Such procedures may include, for example, biopsies,inserting a localized radiation source for direct treatment of acancerous lesion, known as brachytherapy (so as to prevent radiationdamage to tissues near the lesion), injecting a chemotherapy agent intothe cancerous site or removing a cancerous or other lesions.

The aim of all such procedures is to pin-point the target area asprecisely as possible in order to get the most precise biopsy results,preferably from the most active part of a tumor, or to remove such atumor in its entirety, with minimal damage to the surrounding, nonaffected tissues.

This goal is yet to be achieved, as most of the common imagingmodalities such as fluoroscopy, CT, MRI, mammography or ultrasounddemonstrate the position and appearance of the entire lesion withanatomical modifications that the lesion causes to its surroundingtissue, without differentiating between the non-active mass from thephysiologically active part thereof.

Furthermore, prior art radiation emission detectors and (or) biopsyprobes, while being suitable for identifying the location of theradiation site, leave something to be desired from the standpoint offacilitating the removal or other destruction of the detected canceroustissue, with minimal trauma.

The combination of modalities, as is offered by the present invention,can reduce the margin of error in locating such tumors. In addition, thepossibility of demonstrating the position of the active part of a tumorsuperimposed on a scan from an imaging modality that shows the organ ortumor, coupled with the possibility to follow a surgical tool inreference to the afflicted area during a surgical procedure will allowfor a more precise and controlled surgical procedures to take place,minimizing the aforementioned problems.

The present invention addresses these and other issues which are furtherelaborated hereinbelow, and offers the physicians and patients morereliable targeting, which in turn will result in less invasive and lessdestructive surgical procedures and fewer cases of mistaken diagnoses.

SUMMARY OF THE INVENTION

The present invention successfully addresses the shortcomings of thepresently known configurations by providing a radioactive emission probein communication with a position tracking system and the use thereof ina variety of systems and methods of medical imaging and procedures.Specifically, wide-aperture collimation-deconvolution algorithms areprovided, for obtaining a high-efficiency, high resolution image of aradioactivity emitting source, by scanning the radioactivity emittingsource with a probe of a wide-aperture collimator, and at the same time,monitoring the position of the radioactive emission probe, at very finetime intervals, to obtain the equivalence of fine-aperture collimation.The blurring effect of the wide aperture is then correctedmathematically. Furthermore, an imaging method by depth calculations isprovided, based on the attenuation of photons of different energies,which are emitted from the same source, coupled with positionmonitoring.

The present invention has many other applications in the direction oftherapeutics, such as, but not limited to, implanting brachytherapyseeds, ultrasound microwave radio-frequency cryotherapy and localizedradiation ablations.

Implementation of the methods and systems of the present inventioninvolves performing or completing selected tasks or steps manually,automatically, or a combination thereof. Moreover, according to actualinstrumentation and equipment of preferred embodiments of the methodsand systems of the present invention, several selected steps could beimplemented by hardware or by software on any operating system of anyfirmware or a combination thereof. For example, as hardware, selectedsteps of the invention could be implemented as a chip a circuit. Assoftware, selected steps of the invention could be implemented as aplurality of software instructions being executed by a computer usingany suitable algorithms. In any case, selected steps of the method andsystem of the invention could be described as being performed by a dataprocessor, such as a computing platform for executing a plurality ofinstructions.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention is herein described, by way of example only, withreference to the accompanying drawings. With specific reference now tothe drawings in detail, it is stressed that the particulars shown are byway of example and purposes of illustrative discussion of the preferredembodiments of the present invention only, and are presented in thecause of providing what is believed to be the most useful and readilyunderstood description of the principles and conceptual aspects of theinvention. In this regard, no attempt is made to show structural detailsof the invention in more detail than is necessary for a fundamentalunderstanding of the invention, the description taken with the drawingsmaking apparent to those skilled in the art how the several forms of theinvention may be embodied in practice.

In the drawings:

FIG. 1 is a black box diagram of a system according to the teachings ofthe present invention;

FIG. 2 is a perspective view of an articulated arm which serves as aposition tracking system shown carrying a radioactive emission probe inaccordance with the teachings of the present invention;

FIG. 3 is a schematic depiction of a radioactive emission probe carryinga pair of three coaxially aligned accelerometers which serve as aposition tracking system in accordance with the teachings of the presentinvention;

FIG. 4 is a schematic presentation of a radioactive emission probecommunicating with yet another type of a position tracking system inaccordance with the teachings of the present invention;

FIG. 5 is a simplified cross-sectional view of a narrow or wide angleradioactive emission probe used to implement an embodiment of thepresent invention;

FIG. 6 is a presentation of a scanning protocol which can be effectedwith the detector of FIG. 5;

FIG. 7 is a simplified cross-sectional view of a spatially sensitiveradioactive emission probe, e.g., a gamma camera, used to implementanother embodiment of the present invention;

FIG. 8 is a presentation of a scanning protocol which can be effectedwith the detector of FIG. 7;

FIG. 9 demonstrates a system in accordance with the teachings of thepresent invention which employs four position tracking systems forco-tracking the positions of a patient, a radioactive emission probe, animaging modality and a surgical instrument;

FIG. 10 demonstrates the use of a pair of radiation emission detectorsconnected therebetween via a connector, preferably a flexible connectoror a flexible connection to the connector according to the presentinvention;

FIG. 11 is a schematic diagram of a surgical instrument and accompanyingsystem elements according to the teachings of the present invention;

FIG. 12 is a simplified pictorial illustration of an imaging systemconstructed and operative in accordance with a preferred embodiment ofthe present invention, including a radiation probe and position sensor,position tracking system, medical imaging system and coordinateregistration system;

FIG. 13 is a simplified pictorial illustration of a single dimensionimage formation with a nuclear radiation probe attached to a positiontracking system of the system of FIG. 12, in accordance with a preferredembodiment of the present invention;

FIG. 14 is a simplified pictorial plot of detecting a radiation pointsource with the nuclear radiation probe of the system of FIG. 12,without further processing, in accordance with a preferred embodiment ofthe present invention;

FIG. 15 is a simplified flow chart of an averaging algorithm used in theimaging system of FIG. 12, in accordance with a preferred embodiment ofthe present invention;

FIG. 16 is a simplified pictorial plot of detecting a radiation pointsource with the nuclear radiation probe of the system of FIG. 12, withaveraging processing, in accordance with a preferred embodiment of thepresent invention;

FIGS. 17 and 18 are simplified pictorial illustrations of hot cross andhot bar phantom images, respectively, of images produced by a gammaradiation probe of the system of FIG. 12;

FIG. 19 is a simplified flow chart of a minimizing algorithm used in theimaging system of FIG. 12, in accordance with a preferred embodiment ofthe present invention;

FIG. 20 is a simplified pictorial plot of detecting a radiation pointsource with the nuclear radiation probe of the system of FIG. 12, withminimizing processing, in accordance with a preferred embodiment of thepresent invention;

FIG. 21 is a simplified pictorial illustration of an imagereconstruction system constructed and operative in accordance with apreferred embodiment of the present invention, which produces a combinedimage made up of medical images, the position of the peak radiationlocation and the location of a therapeutic instrument;

FIG. 22 is a simplified flow chart of a radiation map reconstructionalgorithm, in accordance with a preferred embodiment of the presentinvention;

FIGS. 23A and 23B are illustrations of radiolabeled patterns observed inimages produced by the system of the invention and by a conventionalgamma camera, respectively, of an autonomous adenoma of a thyroid;

FIGS. 24A and 24B are illustrations of radiolabeled patterns observed inimages produced by the system of the invention and by a conventionalgamma camera, respectively, of suspected Paget's disease of a humerus;

FIGS. 25A and 25B are illustrations of radiolabeled patterns observed inimages produced by the system of the invention and by a conventionalgamma camera, respectively, of chronic osteomyelitis;

FIGS. 26A and 26B are illustrations of radiolabeled patterns observed inimages produced by the system of the invention and by a conventionalgamma camera, respectively, of skeletal metastasis from medulloblastoma;

FIGS. 27A-27I demonstrate the operation of an algorithm provided by thepresent invention for estimating the distribution of radiation sourcesin a control volume;

FIGS. 28A-28F schematically illustrate a handheld probe, in accordancewith preferred embodiments of the present invention;

FIGS. 29A-29B schematically illustrate the manner of calibrating thehandheld probe of FIGS. 28A-28F, in accordance with a preferredembodiment of the present invention;

FIG. 30. schematically illustrates the manner of synchronizing event andposition readings of the handheld probe of FIGS. 28A-28F, in accordancewith a preferred embodiment of the present invention;

FIGS. 31A-31C describe a spatial resolution bar-phantom test of aradioactive emission probe, in accordance with a preferred embodiment ofthe present invention;

FIGS. 32A-32D describe a spatial resolution bar-phantom test of aprior-art probe;

FIGS. 33A-33B illustrate the energy resolution of a single pixel of aradioactive emission probe, in accordance with the present invention;

FIGS. 34A-34C illustrate endoscopic radioactive emission probes, inaccordance with the present invention;

FIG. 35 illustrate a method of calculating the depth of a radiationsource, in accordance with the present invention; and

FIG. 36 illustrate a two-dimensional image of the radioactivity emittingsource, produced by a free-hand scanning of a cancerous prostate gland,ex vivo, in accordance with the present invention.

DESCRIPTION OF THE PREFERRED EMBODIMENTS

The present invention relates to a radioactive emission probe incommunication with a position tracking system and the use thereof in avariety of systems and methods of medical imaging and procedures.Specifically, wide-aperture collimation-deconvolution algorithms areprovided, for obtaining a high-efficiency, high resolution image of aradioactivity emitting source, by scanning the radioactivity emittingsource with a probe of a wide-aperture collimator, and at the same time,monitoring the position of the radioactive emission probe, at very finetime intervals, to obtain the equivalence of fine-aperture collimation.The blurring effect of the wide aperture is then correctedmathematically. Furthermore, an imaging method by depth calculations isprovided, based on the attenuation of photons of different energies,which are emitted from the same source, coupled with positionmonitoring.

The principles and operation of the present invention may be betterunderstood with reference to the drawings and accompanying descriptions.

Before explaining at least one embodiment of the invention in detail, itis to be understood that the invention is not limited in its applicationto the details of construction and the arrangement of the components setforth in the following description or illustrated in the drawings. Theinvention is capable of other embodiments or of being practiced orcarried out in various ways. Also, it is to be understood that thephraseology and terminology employed herein is for the purpose ofdescription and should not be regarded as limiting.

Functional imaging, or the use of radioactive materials to tagphysiologically active tissue within the body of a patient, fordetermining the tissue's localization and demarcation by radioactiveemission probes has been disclosed in the medical literature for atleast forty years. Functional imaging shows the metabolic activity ofbody tissue, since dying or damaged body tissue absorbsradiopharmaceuticals at a different rate from a healthy tissue. Thefunctional image may be used for example, to study cardiac rhythm, orrespiratory rhythm. However, a functional image may not show structural,or anatomic details.

Significant developments in the localization and demarcation of tissuebearing radioactive isotope tags for diagnostic and (or) therapeuticpurposes have occurred since that time. In fact, it is now becoming anestablished practice in the diagnosis and (or) treatment of certaindiseases, e.g., cancer, blood clots, myocardial infarct and abscesses,to introduce monoclonal antibodies or other agents, e.g., fibrinogen,fluorodeoxyglucose labeled with a radioactive isotope (e.g., ⁹⁹ ^(M)Technetium, ⁶⁷Gallium, ²⁰¹Thallium, ¹¹¹Indium, ¹²³Iodine, ¹⁸Fluorine and¹²⁵Iodine) into the body of the patient. Such radiopharmaceuticals tendto localize in particular tissue or cell type, whereas uptake or bindingof the specific radiopharmaceutical is increased in more“physiologically active” tissue such as the active core of a canceroustissue, so that the radiation emitted following nuclear disintegrationsof the isotope can be detected by a radiation detector to betterallocate the active portion of a tumor. Such radiation may be, forexample, α, β⁻, β⁺ and (or) γ radiation.

In another type of applications radioactive substances are used todetermine the level of flow of blood in blood vessels and the level ofperfusion thereof into a tissue, e.g., coronary flow and myocardialperfusion.

Referring now to the drawings, FIG. 1 illustrates a system forcalculating a position of a radioactivity emitting source in asystem-of-coordinates, in accordance with the teachings of the presentinvention, which system is referred to hereinbelow as system 20.

System 20 includes a radioactivity emission detector 22. System 20according to the present invention further includes a position trackingsystem 24. System 24 is connected to and (or) communicating withradioactive emission probe 22 so as to monitor the position of detector22 in a two- or three-dimensional space defined by asystem-of-coordinates 28 in two, three or more, say four, five orpreferably six degrees-of-freedom (X, Y, Z, ρ, θ and φ). System 20further includes a data processor 26. Data processor 26 is designed andconfigured for receiving data inputs from position tracking system 24and from radioactive emission probe 22 and, as is further detailedbelow, for calculating the image of the radioactivity emitting source insystem-of-coordinates 28. As shown in FIG. 10, a pair (or more) ofdetectors 22 connected therebetween via a physical connector, each ofdetectors 22 is position tracked, can be used for calculating the imageof the radioactivity emitting source in system-of-coordinates 28. Ifmore than a single detector 22 is used, detectors 22 are preferablyconnected there between via a connector 29. Connector 29 is preferablyflexible. In the alternative, the connections of detectors 22 toconnector 29 provide the required flexibility.

It will be appreciated that system 20 of radioactivity emission detector22 and position tracking system 24 is inherently different from knownSPECT and PET imaging systems, as well as from other imaging systemssuch as X-ray, Mammography, CT, and MRI, since the motion of detector 22of the present invention is not limited to a predetermined track ortracks, with a respect to an immovable gantry. Rather, detector 22 ofthe present invention is adapted for a variable-course motion, which maybe for example, free-hand scanning, variable-course motion on a linkagesystem, motion within a body lumen, endoscopic motion through a trocarvalve, or another form of variable-course motion.

Position tracking systems per se are well known in the art and may useany one of a plurality of approaches for the determination of positionin a two- or three-dimensional space as is defined by asystem-of-coordinates in two, three and up to six degrees-of-freedom.Some position tracking systems employ movable physical connections andappropriate movement monitoring devices (e.g., potentiometers) to keeptrack of positional changes. Thus, such systems, once zeroed, keep trackof position changes to thereby determine actual positions at all times.One example for such a position tracking system is an articulated arm.

FIG. 2 shows an articulated arm 30 which includes six arm members 32 anda base 34, which can therefore provide positional data in sixdegrees-of-freedom. Monitoring positional changes may be effected in anyone of several different ways. For example, providing each arm member 32with, e.g., potentiometers or optical encoders 38 used to monitor theangle between adjacent arm members 32, to thereby monitor the angularchange of each such arm member with respect to adjacent arm members, soas to determine the position in space of radioactive emission probe 22,which is physically connected to articulated arm 30.

As is shown in FIG. 3 other position tracking systems can be attacheddirectly to radioactive emission probe 22 in order to monitor itsposition in space. An example of such a position tracking system is anassortment of three triaxially (e.g., co-orthogonally) orientedaccelerometers 36 which may be used to monitor the positional changes ofradioactive emission probe 22 with respect to a space. A pair of suchassortments, as is specifically shown in FIG. 3, can be used todetermine the position of detector 22 in six-degrees of freedom.

As is shown in FIGS. 4 and 10, other position tracking systemsre-determine a position irrespective of previous positions, to keeptrack of positional changes. Such systems typically employ an array ofreceivers/transmitters 40 which are spread in known positions in asystem-of-coordinates and transmitter(s)/receiver(s) 42, respectively,which are in physical connection with the object whose position beingmonitored. Time based triangulation and (or) phase shift triangulationare used in such cases to periodically determine the position of themonitored object, radioactive emission probe 22 in this case. Examplesof such a position tracking systems employed in a variety of contextsusing acoustic (e.g., ultrasound) electromagnetic radiation (e.g.,infrared, radio frequency) or magnetic field and optical decoding aredisclosed in, for example, U.S. Pat. Nos. 5,412,619; 6,083,170;6,063,022; 5,954,665; 5,840,025; 5,718,241; 5,713,946; 5,694,945;5,568,809; 5,546,951; 5,480,422 and 5,391,199, which are incorporated byreference as if fully set forth herein.

Radioactive emission probes are well known in the art and may use anyone of a number of approaches for the determination of the amount ofradioactive emission emanating from an object or portion thereof.Depending on the type of radiation, such detectors typically includesubstances which when interacting with radioactive decay emittedparticles emit either electrons or photons in a level which isproportional over a wide linear range of operation to the level ofradiation impinging thereon. The emission of electrons or photons ismeasurable and therefore serves to quantitatively determine radiationlevels. Solid-state detectors in the form of N-type, P-type, PIN-typepixellated or unpixellated include, for example, Ge, Si, CdTe, CdZnTe,CdSe, CdZnSe, HgI₂, TlBrI, GaAs, InI, GaSe, Diamond, TlBr, Pb₂, InP,ZnTe, HgBrI, a-Si, a-Se, BP, GaP, CdS, SiC, AlSb, PbO, BiI₃ and ZnSedetectors. Gas (e.g., CO₂ CH₄) filled detectors include ionizationchamber detectors, proportional chamber detectors and Geiger chamberdetectors. Scintillation detectors include organic scintillator crystalsand liquids, such as C₁₄H₁₀, C₁₄H₁₂, C₁₀H₈, etc., Plastics, NE102A,NE104, NE110, Pilot U and inorganic scintillator crystals, such as NaI,CsI, BGO, LSO, YSO, BaF, ZnS, ZnO, CaWO₄ and CdWO₄. Also known arescintillation fiber detectors. Scintillator coupling includephotomultiplier tube (PMT) of the following types: side-on type, head-ontype, hemispherical type, position sensitive type, icrochannelplate-photomultiplier (MCP-PMTs) and electron multipliers, orphotodiodes (and photodiodes arrays), such as Si photodiodes, Si PINphotodiodes, Si APD, GaAs(P) photodiodes, GaP and CCD.

FIG. 5 shows a narrow angle or wide angle radioactive emission probe22′. Narrow or wide angle radioactive emission probe 22′ includes anarrow slit (collimator) so as to allow only radiation arriving from apredetermined angular direction (e.g., 1°-280° wide angle, preferably1°-80°—narrow angle) to enter the detector. Narrow or wide angleradioactive emission probes especially suitable for the configurationshown in FIG. 10 are manufactured, for example, by Neoprobe, Dublin,Ohio (www.neoprobe.com), USA, Nuclear Fields, USA (www.nufi.com)IntraMedical Imaging, Los Angeles, Calif., USA (www.gammaprobe.com).

As is shown in FIG. 6, such a detector is typically used to measureradioactivity, point by point, by scanning over the surface of aradioactive object from a plurality of directions and distances. In theexample shown, scans from four different directions are employed. Itwill be appreciated that if sufficient radioactivity records arecollected from different angles and distances, and the orientation andposition in space of detector 22′ is simultaneously monitored andrecorded during such scans, a three-dimensional model of a radioactiveregion can be reconstituted and its position in space determined. If twoor more detectors are co-employed, as shown in the configuration of FIG.10, the results may be collected faster.

FIG. 7 shows another example of a radioactive emission probe, aspatially sensitive (pixellated) radioactive emission probe 22″ (such asa gamma camera). Detector 22″, in effect, includes an array of multitudenarrow angle detector units 23. Such an arrangement is used inaccordance with the teachings of the present invention to reduce theamount of measurements and angles necessary to acquire sufficient dataso as to reconstitute a three-dimensional model of the radioactiveobject. Examples of spatially sensitive radioactive emission probesemployed in a variety of contexts are disclosed in, for example, U.S.Pat. Nos. 4,019,057; 4,550,250; 4,831,262; and 5,521,373; which areincorporated by reference as if set forth herein. An additional exampleis the COMPTON detector(http://www.ucl.ac.uk/MedPhvs/posters/giulia/giulia.htm). FIG. 8 shows ascan optionally made by spatially sensitive radioactive emission probe22″ (such as a gamma camera).

A radioactive emission detector of particular advantages for use incontext of the present invention is the Compton gamma probe, since, inthe Compton gamma probe, spatial resolution is independent ofsensitivity and it appears possible to exceed the noise equivalentsensitivity of collimated imaging systems especially for systems withhigh spatial resolution. The Compton probe is a novel type ofgamma-probe that makes use of the kinematics of Compton scattering toconstruct a source image without the aid of mechanical collimators.Compton imaging telescopes were first built in the 1970s forastronomical observations [V. Schoenfelder et al., Astrophysical Journal217 (1977) 306]. The first medical imaging laboratory instrument wasproposed in the early 1980s [M. Singh, Med. Phys. 10 (1983) 421]. Thepotential advantages of the Compton gamma probe include higherefficiency, 3-D imaging without detector motion, and more compact andlightweight system. In the Compton gamma probe, high-energy gamma raysare scattered from a first detector layer (or detectors array) into asecond detector layer array. For each gamma, the deposited energy ismeasured in both detectors. Using a line drawn between these twodetectors, the Compton scattering equation can be solved to determinethe cone of possible direction about this axis on which the gamma raymust have entered the first detector. The intersection of cones frommany events is then developed to locate gamma ray sources in the probe'sfield-of-view. Obviously only coincident events are considered, and themore accurately their energy can be determined, the less uncertaintythere is in the spatial angle of the arrival cone. The probe'selectronic system is combining coincidence measurements across manydetectors and detectors layers with a very good energy resolution. Thechoice of the geometry and the material of the first layer detectorplays a major role in the system imaging capability and depends on (i)material efficiency of single Compton events, in relation to otherinteractions; (ii) detector energy resolution; and (iii) detectorposition resolution. In particular, the overall angular resolutionresults from the combination of two components, related to the energyresolution and to the pixel volume of the detector.

Thus, as now afforded by the present invention, connecting a radioactiveemission probe to a position tracking system, permits simultaneousradioactivity detecting and position tracking at the same time. Thisenables the accurate calculation of the shape, size and contour of theradiating object and its precise position in a system-of-coordinates.

The present invention thus provides a method for defining a position ofa radioactivity emitting source in a system-of-coordinates. The methodis effected by (a) providing a radioactive emission probe which is incommunication with a position tracking system; and (b) monitoringradioactivity emitted from the radioactivity emitting source, while atthe same time, monitoring the position of radioactive emission probe inthe system-of-coordinates, thereby defining the image of theradioactivity emitting source in the system-of-coordinates.

It will be appreciated by one of skills in the art that the modelproduced by system 20 is projectable onto any of the othersystems-of-coordinates, or alternatively, the system-of-coordinatesdefined by position tracking system 24 may be shared by other positiontracking systems, as is further detailed hereinbelow, such that no suchprojection is required.

Thus, as is further shown in FIG. 1, system 20 of the present inventioncan be used for calculating a position of a radioactivity emittingsource in a first system-of-coordinates 28 and further for projectingthe image of the radioactivity emitting source onto a secondsystem-of-coordinates 28′. The system includes radioactive emissionprobe 22, position tracking system 24 which is connected to and (or)communicating with radioactive emission probe 22, and data processor 26which is designed and configured for (i) receiving data inputs fromposition tracking system 24 and from radioactive emission probe 22; (ii)calculating the image of the radioactivity emitting source in the firstsystem-of-coordinates; and (iii) projecting the image of theradioactivity emitting source onto the second system-of-coordinates.

A method for calculating a position of a radioactivity emitting sourcein a first system-of-coordinates and for projecting the image of theradioactivity emitting source onto a second system-of-coordinates isalso offered by the present invention. This method is effected by (a)providing a radioactive emission probe being in communication with aposition tracking system; and (b) monitoring radioactivity being emittedfrom the radioactivity emitting source, while at the same time,monitoring the position of the radioactive emission probe in the firstsystem-of-coordinates, thereby defining the image of the radioactivityemitting source in the first system-of-coordinates and projecting theimage of the radioactivity emitting source onto the secondsystem-of-coordinates.

It will be appreciated that the combination of a radioactive emissionprobe and a position tracking system connected thereto and (or)communicating therewith allows a suitable data processor to generate atwo- or three-dimensional image of the radioactivity emitting source. Analgorithm can be used to calculate image intensity based on, forexample, a probability function which averages radiation counts andgenerates an image in which the shorter the time interval betweenradioactive counts, the brighter the image and vise versa, whiledown-compensating when a location is re-scanned. A free-hand scanningwith a directional detector can be employed for this purpose.

In one embodiment, when scanning a body area with the detector, thedetector is made to follow a three-dimensional surface, which definesthe body curvature and in effect is used also as a position trackingpointer. This information can be used to define the position of theradioactive source with respect to the outer surface of the body, so asto create a three-dimensional map of both the radioactive source and ofthe body curvature. This approach can also be undertaken in opensurgeries, such as open chest surgeries so as to provide the surgeon inreal time with information concerning the functionality of a tissue.

The radioactive emission probe, which can be used in context of thepresent invention can be a beta emission detector, a gamma emissiondetector, a positron emission detector or any combination thereof. Adetector that is sensitive to both beta and (or) positron and gammaemission can be used to improve localization by sensing for examplegamma emission distant from the source and sensing beta or positronsemission closer to the source. A beta detector is dedicated for thedetection of either electrons from sources such as ¹³¹Iodine, orpositrons from sources such as ¹⁸Fluorine. A gamma detector can bedesigned as a single energy detector or as a detector that candistinguish between different types of energies, using the lightintensity in the scintillator as a relative measure of the gamma energy.Also, the detector can be designed to utilize coincidence detection byusing detectors facing one another (180 degrees) with the examined organor tissue in-between. The radiation detector can have differentcollimators with different diameters. A large bore will be used for highsensitivity with lower resolution while a small bore collimator willhave higher resolution at the expense of lower sensitivity.

Another possibility is to have a the collimator moving or rotating withthe opening eccentric so that a different solid angle is exposed to theincoming photons at any one time, thus gathering the photons fromoverlapping volumes at different time intervals. The rest of the imageprocessing is similar if the probe moves or if the collimator eccentricopening moves.

System 20 of the present invention can be used in concert with othermedical devices, such as, but not limited to, any one of a variety ofimaging modalities and (or) surgical instruments.

Structural imaging modalities, which provide anatomic, or structuralmaps of the body, are well known in the art. The main modalities thatserve for two-(projectional or cross sectional) or three- (cosequtivecross sectional) dimensional imaging are a planer X-ray imager, afluoroscope, a computerized tomography scanner, a magnetic resonanceimager an ultrasound imager, an impedance imager, and an optical camera.

Medical images taken of the human body are typically acquired ordisplayed in three main orientations (i) coronal orientation: in a crosssection (plane), for example, across the shoulders, dividing the bodyinto front and back halves; (ii) sagittal orientation: in a crosssection (plane), for example, down the middle, dividing the body intoleft and right halves; and (iii) axial orientation: in a cross section(plane), perpendicular to the long axis of the body, dividing the bodyinto upper and lower halves. Oblique views can also be acquired anddisplayed.

Various types of X-ray imaging are central to diagnosis of many types ofcancer. Conventional X-ray imaging has evolved over the past 100 years,but the basic principal is still the same as in 1895, when firstintroduced. An X-ray source is turned on and X-rays are radiated throughthe body part of interest and onto a film cassette positioned under orbehind the body part. The energy and wavelength of the X-rays allowsthem to pass through the body part and create the image of the internalstructures like bones. As the X-rays pass through the hand, forinstance, they are attenuated by the different density tissues theyencounter. Bone attenuates a great deal more of the X-rays than the softtissue surrounding it because of its grater density. It is thesedifferences in absorption and the corresponding varying exposure levelof the film that creates the images. In fact, X-ray imaging results in aprojection of the integrated density of column-voxels defined by theX-rays as they pass through the body.

Fluoroscopy is a method based on the principals of film X-ray that isuseful for detecting disorders and tumors in the upper gastro-intestinal(GI) system (for example, the stomach and intestines). Fluoroscopicimaging yields a moving X-ray picture. The physician can watch thescreen and see an image of the patient's body (for example the beatingheart). Fluoroscopic technology improved greatly with the addition oftelevision cameras and fluoroscopic “image intensifiers”. Today, manyconventional X-ray systems have the ability to switch back and forthbetween the radiographic and fluoroscopic modes. The latest X-raysystems have the ability to acquire the radiograph or fluoroscopic movieusing digital acquisition.

Computed Tomography (CT) is based on the X-ray principal, where the filmis replaced by a detector that measures the X-ray profile. Inside thecovers of the CT scanner is a rotating frame which has an X-ray tubemounted on one side and the detector mounted on the opposite side. A fanbeam of X-ray is created as the rotating frame spins the X-ray tube anddetector around the patient. Each time the X-ray tube and detector makea 360° rotation, an image or “slice” has been acquired. This “slice” iscollimated to a thickness between 1 mm and 10 mm using lead shutters infront of the X-ray tube and X-ray detector.

As the X-ray tube and detector make this 360° rotation, the detectortakes numerous profiles of the attenuated X-ray beam. Typically, in one360° lap, about 1,000 profiles are sampled. Each profile is subdividedspatially by the detectors and fed into about 700 individual channels.Each profile is then backwards reconstructed (or “back projected”) by adedicated computer into a two-dimensional image of the “slice” that wasscanned.

The CT gantry and table have multiple microprocessors that control therotation of the gantry, movement of the table (up/down and in/out),tilting of the gantry for angled images, and other functions such asturning the X-ray beam on an off. The CT contains a slip ring thatallows electric power to be transferred from a stationary power sourceonto the continuously rotating gantry. The innovation of the power slipring has created a renaissance in CT called spiral or helical scanning.These spiral CT scanners can now image entire anatomic regions like thelungs in a quick 20 to 30 second breath hold. Instead of acquiring astack of individual slices which may be misaligned due to slight patientmotion or breathing (and lung/abdomen motion) in between each sliceacquisition, spiral CT acquires a volume of data with the patientanatomy all in one position. This volume data set can then becomputer-reconstructed to provide three-dimensional models such as ofcomplex blood vessels like the renal arteries or aorta. Spiral CT allowsthe acquisition of CT data that is perfectly suited forthree-dimensional reconstruction.

MR Imaging is superior to CT in detecting soft tissue lesions such astumors as it has excellent contrast resolution, meaning it can showsubtle soft-tissue changes with exceptional clarity. Thus, MR is oftenthe method of choice for diagnosing tumors and for searching formetastases. MR uses magnetic energy and radio waves to create single orconsequtive cross-sectional images or “slices” of the human body. Themain component of most MR systems is a large tube shaped or cylindricalmagnet. Also, there are MR systems with a C-shaped magnet or other typeof open designs. The strength of the MR systems magnetic field ismeasured in metric units called “Tesla”. Most of the cylindrical magnetshave a strength between 0.5 and 1.5 Tesla and most of the open orC-shaped magnets have a magnetic strength between 0.01 and 0.35 Tesla.

Inside the MR system a magnetic field is created. Each total MRexamination typically is comprised of a series of 2 to 6 sequences. An“MR sequence” is an acquisition of data that yields a specific imageorientation and a specific type of image appearance or “contrast”.During the examination, a radio signal is turned on and off, andsubsequently the energy which is absorbed by different atoms in the bodyis echoed or reflected back out of the body. These echoes arecontinuously measured by “gradient coils” that are switched on and offto measure the MR signal reflecting back. In the rotating frame ofreference, the net magnetization vector rotate from a longitudinalposition a distance proportional to the time length of the radiofrequency pulse. After a certain length of time, the net magnetizationvector rotates 90 degrees and lies in the transverse or x-y plane. It isin this position that the net magnetization can be detected on MRI. Theangle that the net magnetization vector rotates is commonly called the‘flip’ or ‘tip’ angle. At angles greater than or less than 90 degreesthere will still be a small component of the magnetization that will bein the x-y plane, and therefore be detected. Radio frequency coils arethe “antenna” of the MRI system that broadcasts the RF signal to thepatient and (or) receives the return signal. RF coils can bereceive-only, in which case the body coil is used as a transmitter; ortransmit and receive (transceiver). Surface coils are the simplestdesign of coil. They are simply a loop of wire, either circular orrectangular, that is placed over the region of interest.

A digital computer reconstructs these echoes into images of the body. Abenefit of MRI is that it can easily acquire direct views of the body inalmost any orientation, while CT scanners typically acquirecross-sectional images perpendicular or nearly perpendicular to the longbody axis.

Ultrasound imaging is a versatile scanning technique that uses soundwaves to create images of organs or anatomical structures in order tomake a diagnosis. The ultrasound process involves placing a small devicecalled a transducer, against the skin of the patient near the region ofinterest, for example, against the back to image the kidneys. Theultrasound transducer combines functions of emitting and receivingsound. This transducer produces a stream of inaudible, high frequencysound waves which penetrate into the body and echo off the organsinside. The transducer detects sound waves as they echo back from theinternal structures and contours of the organs. Different tissuesreflect these sound waves differently, causing a signature which can bemeasured and transformed into an image. These waves are received by theultrasound machine and turned into live pictures with the use ofcomputers and reconstruction software.

Ultrasound scanning has many uses, including: diagnosis of disease andstructural abnormalities, helping to conduct other diagnosticprocedures, such as needle biopsies etc.

There are limitations to some ultrasound techniques: good images may notbe obtained in every case, and the scan may not produce as preciseresults as some other diagnostic imaging procedures. In addition, scanresults may be affected by physical abnormalities, chronic disease,excessive movement, or incorrect transducer placement.

Both two- (cross sectional) and three- (consequtive cross-sectional)ultrasound imaging techniques are available nowadays. Worth mentioningis the Dopler three-dimensional ultrasound imaging.

In many cases imaging modalities either inherently include (e.g.,fluoroscope, CT, MRI) and (or) are integrated withposition-tracking-systems, which enable the use of such systems toreconstruct three-dimensional image models and provide their position ina system-of-coordinates.

It will be appreciated that, similar to the vision system, also anoptical camera can be used to generate three-dimensional imagery dateaccording to the present invention by imaging a body from a plurality(at least two) directions. This type of imaging is especially applicablein open chest surgeries or other open surgeries. Software forcalculating a three-dimensional image from a pair of stereoscopic imagesis well known in the art.

Thus, as used herein and in the claims section that follows, the phrase“three-dimensional imaging modality” refers to any type of imagingequipment which includes software and hardware for generating athree-dimensional image. Such an equipment can generate athree-dimensional image by imaging successive cross-sections of a body,e.g., as if viewed from a single direction. Alternatively, such anequipment can generate a three-dimensional image by imaging a body fromdifferent angles or directions (typically two angles) and thereaftercombining the data into a three-dimensional image.

In accordance with the present invention, a structural imaging probe,for example, an ultrasound probe, may be combined with a positiontracking system, in a manner analogous to System 20 of radioactivityemission detector 22 and position tracking system 24 (FIG. 1).Similarly, a miniature MRI probe, for example, as taught by U.S. Pat.No. 5,572,132, to Pulyer, et al., entitled, “MRI probe for externalimaging,” whose disclosure is incorporated herein by reference, whichteaches an MRI catheter for endoscopical imaging of tissue of the arterywall, rectum, urinal tract, intestine, esophagus, nasal passages, vaginaand other biomedical applications, may be used.

Additionally, a structural imaging probe, for example, an ultrasoundprobe, may be combined with system 20 of radioactivity emission detector22 and position tracking system 24 (FIG. 1), so as to have:

i. structural imaging;

ii. functional imaging; and

iii. position tracking,

simultaneously. Similarly, other structural imaging modalities, asknown, may be combined with system 20 of FIG. 1.

Furthermore, the structural imaging modality may have an additional,dedicated position tracking system, or share position tracking system 24of radioactivity emission detector 22.

Position tracking may also be accomplished by using imaging informationfrom the structural imaging system. This may be accomplished by trackingrelative changes from one image to another to determine relative motion.

Surgical instruments are also well known in the art and may use any oneof a plurality of configurations in order to perform minimally-invasivesurgical procedures. Examples include laser probes, cardiac andangioplastic catheters, endoscopic probes, biopsy needles, aspirationtubes or needles, resecting devices, ultrasonic probes, fiber opticscopes, laparoscopy probes, thermal probes and suction/irrigationprobes. Examples of such surgical instruments employed in a variety ofmedical contexts are disclosed in, for example, U.S. Pat. Nos.6,083,170; 6,063,022; 5,954,665; 5,840,025; 5,718,241; 5,713,946;5,694,945; 5,568,809; 5,546,951; 5,480,422 5,391,199, 5,800,414;5,843,017; 6,086,554; 5,766,234; 5,868,739; 5,911,719; 5,993,408;6,007,497; 6,021,341; 6,066,151; 6,071,281; 6,083,166 and 5,746,738,which are incorporated by reference as if fully set forth herein.

For some applications, examples of which are provided in the list ofpatents above, surgical instruments are integrated withposition-tracking-systems, which enable to monitor the position of suchinstruments while placed in and guided through the body of a treatedpatient.

According to a preferred embodiment of the present invention, thesurgical instrument is equipped with an additional radioactive emissionprobe attached thereto or placed therein. This additional detector isused, according to preferred embodiments of the invention, to fine tunethe location of radioactive emission from within the body, and in closerproximity to the radioactive source. Since the surgical tool ispreferably in communication with a position-tracking system, theposition of the additional detector can be monitored and its readoutsused to fine tune the position of the radioactive source within thebody. Thus, according to this aspect of the present invention, at leastone extracorporeal detector and an intracorporeal detector are used inconcert to determine the position of a radioactive source in the body inhighest precision. The extracorporeal detector provides the generalposition of the source and is used for directing the surgical instrumentthereto, whereas the intracorporeal detector is used for reassuringprior to application of treatment or retrieval of biopsy that indeed thesource was correctly targeted at the highest precision.

While according to a presently preferred embodiment of the invention twodetectors, one extracorporeal and one intracorporeal, are employed asdescribed above, for some applications a single intracorporeal detectormay be employed, which detector is attached to or integrated with asurgical instrument whose position is tracked.

The use of intracorporeal and extracorporeal detectors calls for carefulchoice of the radioactive isotope employed with the radiopharmaceutical.While the extracorporeal detector can be constructed with a suitablecollimator for handling strong radiation, such as gamma radiation, theintracorporeal detector is miniature by nature and is limited in designand construction by the construction of the surgical instrument withwhich it is employed. Since collimators for high energy (80-511 KeV)gamma radiation are robust in nature, they are not readily engageablewith miniature detectors. Electron (beta) and positron radiation arecharacterized by: (i) they highly absorbed by biological tissue as theyare of lower energy and higher chemical reactivity; and (ii) they arereadily collimated and focused by thin metal collimators. It is alsopossible to use low energy gamma radiation (10-30 KeV) forintracorporeal applications since the collimation of these gamma photonscan be achieved with thin layers of Tantalum or Tungsten. As such, theradio pharmaceutical of choice is selected to emit both gamma and betaand (or) positron radiation, whereas the extracorporeal detector is setto detect the high energy gamma radiation, whereas the intracorporealdetector is set to detect the low energy gamma, beta and (or) positronradiation. Isotopes that emit both high energy gamma and (or) low energygamma, beta and (or) positron radiation and which can be used per se oras a part of a compound as radiopharmaceuticals include, withoutlimitation, ¹⁸F, ¹¹ In and ¹²³I in radiopharmaceuticals, such as, butnot limited to, 2-[¹⁸F]fluoro-2-deoxy-D-glucose (¹⁸FDG),¹¹¹In-Pentetreotide ([¹¹¹In-DTPA-D-Phe¹]-octreotide),L-3-[¹²³I]-Iodo-alpha-methyl-tyrosine (IMT),O-(2-[¹⁸F]fluoroethyl)-L-tyrosine (L-[¹⁸F]FET), ¹¹¹In-Capromab Pendetide(CYT-356, Prostascint) and ¹¹¹In-Satumomab Pendetide (Oncoscint).

FIG. 11 illustrates a system in accordance with this aspect of thepresent invention. A surgical instrument 100 is shown connected to aresection/aspiration control element 102 as well known in the art.Surgical instrument 100 includes a radioactive emission probe 104, whichhas a collimator 106 for collimating low energy gamma, beta and (or)positron radiation. In some embodiments, as indicated by arrow 108,detector 104 may be translated within instrument 100. A positiontracking system having one element thereof 110 attached to instrument100 and another element thereof 112 at a fixed location serves tomonitor the position of instrument 100 at all times in two, three and upto six degrees of freedom. Radioactive emission probe 104 communicateswith a counter 114 for counting low energy gamma, beta and (or) positronradiation. All the data is communicated to, and processed by, aprocessor 116. The 2D or 3D data may be projected and displayed alongwith 2D or 3D imaging data derived from an imaging modality using ashared presentation device as described elsewhere herein. A real orvirtual image of the surgical instrument itself may also beco-displayed. Examples of commercially available radiation emissiondetectors that can fit inside, for example, a biopsy needle includescintillating plastic optical fibers like S101 and S104, manufactured byPPLASTIFO or an optical fiber communicating with a scintillator (eitherdetector paint or scintillation crystal) at the fiber edge. The level ofdetected radiation can be reported visually or by an audio signal, as iswell known in the art.

Thus, a surgical instrument equipped with a radiation emission detectorand which is connected to and (or) communicating with a positiontracking system forms one embodiment of this aspect of the presentinvention. Such a design acting in concert with either conventionalimaging modalities and (or) extracorporeal radiation emission detectorsform other embodiments of this aspect of the invention. In all cases, asurgical instrument equipped with a radiation emission detector andwhich is connected to and (or) communicating with a position trackingsystem serves for in situ fine tuning of a radioactive source in thebody.

It will be appreciated that in some minimally-invasive procedures eventhe position of the patient him or herself is monitored via a positiontracking system, using, for example, electronic or physical fiducialmarkers attached at certain locations to the patient's body.

Thus, as is further detailed hereinbelow, by projecting thethree-dimensional data and positions received from any of the abovementioned devices into a common system of coordinates, or alternatively,employing a common position tracking system for all of these devices,one can integrate the data into a far superior and comprehensivepresentation.

An example to this desired outcome is shown in FIG. 9. In the embodimentshown, four independent position tracking systems 50, 52, 54 and 56 areused to track the positions of a patient 58, an imaging modality 60, aradioactive emission probe 62 and a surgical instrument 64 in fourindependent systems-of-coordinates 66, 68, 70 and 72, respectively. Ifthe patient is still, no tracking of the patient's position is required.

It will be appreciated that any subset or all of the position trackingsystems employed may be integrated into one or more common positiontracking systems, and (or) that any subset or all of the positiontracking systems employed may share one or more systems-of-coordinates,and further that any positional data obtained by any of the positiontracking systems described in any of the systems-of coordinates may beprojected to any other system of coordinates or to an independent(fifth) system of coordinates 74. In one preferred embodiment,applicable for applications at the torso of the patient, the system ofcoordinates is a dynamic system of coordinates which takes into accountthe chest breathing movements of the patient during the procedure.

As indicated at 76, the raw data collected by detector 62 is recordedand, as indicated at 78, the position and the radioactive data recordsare used to generate a three-dimensional model of a radiopharmaceuticaluptaking portion of a body component of the patient.

Similarly, as indicated at 80, the imagery data collected by imagingmodality 60 is recorded and the position and the imagery data recordsare used to generate a three-dimensional model of the imaged bodycomponent of the patient.

All the data collected is then fed into a data processor 82 whichprocesses the data and, as indicated at 84, generates a combined orsuperimposed presentation of the radioactive data and the imagery data,which is in positional context with patient 58 and surgical instrument64.

Instrument 64, which by itself can be presented in context of thecombined presentation, may then be used to perform the procedure mostaccurately. Processor 82 may be a single entity or may include aplurality of data processing stations which directly communicate with,or even integral to, any one or more of the devices described.

Additionally or alternatively, a structural imaging probe, for example,an ultrasound probe, or another structural probe, as known, may beincorporated with the surgical instrument.

The present invention provides a major advantage over prior art designsbecause it positionally integrates data pertaining to a body portion asretrieved by two independent imaging techniques, conventional imagingand radioactive imaging, to thereby provide a surgeon with the abilitythe fine point the portion of the body to be sampled or treated.

It will be appreciated that subsets of the devices described in FIG. 9may be used as stand-alone systems. For example, a combination ofdetector 62 with its position-tracking system and instrument 64 with itsposition-tracking-system may in some instances be sufficient to performintrabody procedures. For mere diagnostic purposes, without biopsy, acombination of detector 62 position-tracking-system and modality 60position-tracking-system are sufficient.

Reference is now made to FIG. 12, which illustrates an imaging system200 constructed and operative in accordance with a preferred embodimentof the present invention. Imaging system 200 preferably includes aradiation probe 202, such as the narrow angle radioactive emission probe22′ described hereinabove with reference to FIGS. 5 and 10.

A position sensor 204 is provided for sensing the position of radiationprobe 202. Position sensor 204 may be physically attached to radiationprobe 202, or may be distanced therefrom. Position sensor 204 transmitsthe sensed position data to a position tracking system 206. Positiontracking system 206 may be a system like position tracking system 24,described hereinabove with reference to FIG. 1, and position sensor 204may be any kind of sensor applicable for such position tracking systems.

Another method which can be used to locate the source of radiationemission is by using a small hand held gamma camera 205 (such as theDigiRad 2020tc Imager TM, 9350 Trade Place, San Diego, Calif.92126-6334, USA), attached to position sensor 204.

Position tracking system 206 enables radiation probe 202 to freely scanback and forth in two- or three-dimensions over the area of interest ofthe patient, preferably incrementing a short distance between each scanpass. Position tracking system 206 tracks the position of radiationprobe 202 with respect to a position tracking coordinate system, such asX_(p), Y_(p) and Z_(p), with an origin O_(p).

Imaging system 200 also includes a medical imaging system 208, such as,but not limited to, computed or computerized tomography (CT), magneticresonance imaging (MRI), ultrasound imaging, positron emissiontomography (PET) and single photon emission computed tomography (SPECT),for example. Medical imaging system 208 provides images of a patient 209with respect to a medical imaging coordinate system, such as X_(m),Y_(m) and Z_(m), with an origin O_(m).

Imaging system 200 also includes a coordinate registration system 210,such as that described in commonly owned, now abandoned, U.S. patentapplication Ser. No. 09/610,490, entitled, “method for registeringcoordinate systems,” the disclosure of which is incorporated herein byreference. Coordinate registration system 210 is adapted to register thecoordinates of the position tracking coordinate system with those of themedical imaging coordinate system.

Position tracking system 206, medical imaging system 208 and coordinateregistration system 210 are preferably in wired or wirelesscommunication with a processing unit 212 (also referred to as a dataprocessor 212).

In operation of imaging system 200, after administration of aradiopharmaceutical to patient 209, a clinician/physician/surgeon (notshown) may move or scan radiation probe 202 about a target area underexamination. A physiological activity map of the target area is obtainedby measuring the radiation count rate with radiation probe 202, and bycorrelating the count rate with the count rate direction with positiontracking system 206, which follows the motion of the moving or scanningradiation probe 202.

Reference is now made to FIG. 13, which illustrates image formation withradiation probe 202, in accordance with a preferred embodiment of thepresent invention. For the purposes of simplicity, the example shown inFIG. 13 is for a single dimension image formation, but it is readilyunderstood that the same principles hold true for any other dimensionalimage formation.

In one example of carrying out the invention, radiation probe 202 may bea gamma ray detector probe that comprises a collimator 211 and radiationdetector 213. Gamma rays are projected through the probe collimator 211onto radiation detector 213, which produces electronic signals inaccordance with the radiation detected. Radiation probe 202 sends pulsesto a probe counter 215 which may include a pulse height analyzer circuit(not shown). The pulse height analyzer circuit analyzes the electronicsignals produced by radiation detector 213. If the electronic signalsare within a selected energy window, the level of radiation, i.e.,number of radiation counts, is counted by probe counter 215.

Examples of suitable radiation detectors include a solid state detector(SSD) (CdZnTe, CdTe, HgI, Si, Ge, and the like), a scintillationdetector (NaI(Tl), LSO, GSO, CsI, CaF, and the like), a gas detector, ora scintillating fiber detector (S101, S104, and the like).

The position sensor 204 associated with the radiation probe 202 sensesthe position of radiation probe 202, and position tracking system 206calculates and monitors the motion of radiation probe 202 with respectto the position tracking coordinate system. The motion is calculated andmonitored in two, three and up to six dimensions—the linear directionsof the X, Y and Z axes as well as rotations about the X, Y and Z axes,i.e., rotational angles ρ, θ and φ, respectively.

Examples of suitable position tracking systems include a measurementmechanical arm (FaroArm, http://www.faro.com/products/faroarm.asp),optical tracking systems (Northern Digital Inc., Ontario, CanadaNDI-POLARIS passive or active systems), magnetic tracking systems(NDI-AURORA), infrared tracking systems (E-PEN system,http://www.e-pen.com), and ultrasonic tracking systems (E-PEN system),for example.

Processing unit 212 combines the radiation probe count rate from probecounter 215 together with the positional information from positiontracking system 206, and uses an imaging software algorithm 217 to forma two-dimensional or three-dimensional radiotracer-spread image of thetarget area inside the patient's body. The spatial probe positionstogether with the spatial count rates may be stored in memory ordisplayed on a computer monitor 214 as a pattern of marks correspondingto the spatial and count rate position.

An example of such a pattern is shown in FIG. 14, which illustrates asingle-dimensional, unprocessed simulation of a radiation point source218 (FIG. 13), 30 mm deep inside the human body, detected by using a 10mm nuclear radiation probe 202 coupled to position tracking system 206.The graph of FIG. 14 indicates to a physician that there is a peak countrate of about 500 in the probe position of about 50 mm.

In one embodiment of the invention, the imaging software algorithm 217employs an averaging process to refine the curve of FIG. 14. Thisaveraging process will now be described with reference to FIG. 15.

Probe counter 215 feeds probe count rate information N(Xc, Yc, Zc, ρ, θ,φ to processing unit 212 (step 301). Position sensor 204 feeds probeposition information (Xc, Yc, Zc, ρ, θ, φ) to processing unit 212 (step302). Probe parameters (such as its physical size, dx, dy, dz) are alsoinput into processing unit 212 (step 303).

Processing unit 212 then finds all the voxels (i.e., volume pixels) thatrepresent the probe volume in the processing unit memory (step 304),i.e., Xc+dx, Yc+dy, Zc+dz. Processing unit 212 calculates the number oftimes that the calculation process has been done in each voxel from thebeginning of the image formation (step 305), i.e., M(Xc+dx, Yc+dy,Zc+dz). Processing unit 212 then calculates the new average count ratevalues in each voxel (step 306), in accordance with the formula:

N(Xc+dx, Yc+dy, Zc+dz)=[N(Xc+dx, Yc+dy, Zc+dz)+N(Xc, Yc, Zc, ρ, θ,φ)]/[M (Xc+dx, Yc+dy, Zc+dz)+1]

Processing unit 212 then corrects the display image that represents theperceived voxels at N(Xc+dx, Yc+dy, Zc+dz) (step 307). The algorithmthen repeats itself for the next probe position (step 308).

The resulting graph of the averaging algorithm of FIG. 15, as applied tothe example of FIG. 14, is shown in FIG. 16.

FIGS. 17 and 18 respectively show examples of a hot cross phantom imageand a hot 4.77 mm bar phantom image, produced by a gamma radiation probecoupled with position tracking system 206 and the averaging algorithm ofFIG. 15. The probe images were made by using a probe, EG&G Ortec NaI(Tl)model 905-1 (thickness=1″, diameter=1″) connected to a ScintiPack model296. The position tracking system used was the Ascension miniBIRD,commercially available from Ascension Technology Corporation, P.O. Box527, Burlington, Vt. 05402 USA(http://www.ascension-tech.com/graphic.htm). The magnetic tracking andlocation systems of Ascension Technology Corporation use DC magneticfields to overcome blocking and distortion from nearby conductivemetals. Signals pass through the human body without attenuation.

In another embodiment of the invention, the imaging software algorithm217 may employ a minimizing process to refine the curve of FIG. 14 as isnow described with reference to FIG. 19.

Probe counter 215 feeds probe count rate information N(Xc, Yc, Zc, ρ, θ,φ) to processing unit 212 (step 401). Position sensor 204 feeds probeposition information (Xc, Yc, Zc, ρ, θ, φ) to processing unit 212 (step402). Probe parameters (such as its physical size, dx, dy, dz) are alsoinput into processing unit 212 (step 403).

Processing unit 212 then finds all the voxels that represent the probevolume in the processing unit memory (step 404), i.e., Xc+dx, Yc+dy,Zc+dz. From the voxels that represent the probe volume in the processingunit memory, processing unit 212 finds those that have a higher countrate value N(Xc+dx, Yc+dy, Zc+dz) than the inputted probe count rateN(Xc, Yc, Zc, ρ, θ, φ) (step 405). Processing unit 212 then changes thehigher count rate voxels to that of inputted probe count rate N(Xc, Yc,Zc, ρ, θ, φ) (step 406), and corrects the display image at the highercount rate voxels N(Xc+dx, Yc+dy, Zc+dz) (step 407). The algorithm thenrepeats itself for the next probe position (step 408).

The resulting graph of the minimizing algorithm of FIG. 19, as appliedto the example of FIG. 14, is shown in FIG. 20.

Reference is now made to FIG. 21, which illustrates an imagereconstruction system 450, constructed and operative in accordance witha preferred embodiment of the present invention. Image reconstructionsystem 450 produces a combined image 451 made up of the images comingfrom the medical imaging system 208 with the position of the peakradiation location (and its uncertainty area) from processing unit 212,together with the location of a therapeutic instrument 452, such as abiopsy needle. The combined image 451 allows the physician to betterassess the relative position of therapeutic instrument 452 in relationto the anatomical image (from medical imaging system 208) and theposition of the radioactive area as inferred by the radiation detectionalgorithm.

In accordance with preferred embodiments of the present invention, imageacquisition and reconstruction algorithms are provided, for imageacquisition with wide-aperture collimation and image reconstructionwhich includes deconvolution, for resolution enhancement. The overallalgorithms are herethereto referred to as wide-aperturecollimation-deconvolution algorithms.

In essence, the wide-aperture collimation-deconvolution algorithmsenable one to obtain a high-efficiency, high resolution image of aradioactivity emitting source, by scanning the radioactivity emittingsource with a probe of a wide-aperture collimator, and at the same time,monitoring the position of the radioactive emission probe, at very finetime intervals, to obtain the equivalence of fine-aperture collimation.The blurring effect of the wide aperture is then correctedmathematically.

The wide-aperture collimation-deconvolution algorithms are describedhereinbelow, in conjunction with FIGS. 27A-27I, and FIG. 22.Experimental results, showing images produced by the wide-aperturecollimation-deconvolution algorithms of the present invention, incomparison to a conventional gamma camera are seen in FIGS. 23A-26B, andFIG. 36, hereinbelow.

Referring further to the drawings, FIGS. 27A-27I describe wide-aperturecollimation-deconvolution algorithms, in accordance with preferredembodiments of the present invention. The wide-aperturecollimation-deconvolution algorithms are designed for estimating thedistribution of radiation sources in a control volume, thus constructingan image of the radiation sources in the control volume. For simplicity,it is assumed that the radiation sources comprise dot sources thatradiate uniformly in all directions, are localized, and are smoothlydistributed in the control volume.

Consider a single-pixel, wide-bore-collimator probe 707, seen in FIG.27H, formed for example, as a lead tube collimator 708, having a length,L_(collimator) of about 20 mm and a radiation detector 706, for example,a solid state CdZnTl crystal, at its base, with a diameter, D_(detector)of about 10 mm, so that the ratio of L_(collimator) to D_(detector) isabout 2. An angular opening δ of wide-bore-collimator probe 707 may beabout 40 degrees; wide-bore-collimator probe 707 is sensitive to anyincident photons within the confining of about a 40-degree angularopening. It will be appreciated that single-pixel, wide-bore-collimatorprobes of other dimensions and ratios of L_(collimator) to D_(detector)are similarly possible. For example, the ratio of L_(collimator) toD_(detector) may be about 1, or about 3.

By contrast, as seen in FIG. 27I, conventional gamma cameras have amulti-pixel, small-aperture collimation probe, wherein acollimator-pixel arrangement 709 is about 2 mm in diameter and 40 mm inlength, so that the ratio of L_(collimator) to D_(detector) is about 20.An angular opening c of the small-aperture collimation probe is verysmall. In consequence, wide-bore-collimator probe 707 may be about 30times more sensitive than the small-aperture collimation probe ofcollimator—pixel arrangement 709.

Yet, there are disadvantages to wide-bore-collimator probe 707:

-   -   1. a single-pixel probe does not lend itself to the generation        of an image; and    -   2. a wide aperture blurs the information regarding the direction        from which the radiation comes.

With regard to the first point, the generation of an image, aswide-bore-collimator probe 707 is moved in space, while its position andangular orientation are accurately tracked, at very short timeintervals, for example, of about 100 or 200 ms, or any other short timeinterval, a one-, two-, or three-dimensional image of counting rate as afunction of position can be obtained, by data processing.

More specifically, knowing the position and orientation of probe 707 ateach time interval and the photon count rate at that position andorientation can be used to reconstruct the radiation density of anunknown object, even in three dimensions.

With regard to the second point, the blurring effect of the wide-angleaperture, while it is known that the information is blurred, orconvolved, by the wide-aperture collimator, a deconvolution process maybe used to obtain dependable results. Moreover, the convolving functionfor a wide aperture depends only on the geometry of the collimator andmay be expressed as a set of linear equations that can be readilysolved.

Thus, in accordance with the present invention, a wide-bore collimatorprobe, having a position tracking system, operating at very short timeintervals, and connected to a data processor for wide-aperturecollimation-deconvolution algorithms, can be used for the constructionof an image of radiation density in one-, two-, or three-dimensions,with high degree of accuracy.

It will be appreciated that similar analyses may be employed for wideaperture probes having non-circular collimators. For example, thewide-aperture probe may have a rectangular collimator of a single cellor of a coarse grid of cells. Additionally, the collimator may be awide-angle collimator. A corresponding deconvolution function may beused by the data processor.

Reference is now made to FIGS. 27A and 27B, which illustrate a radiationsensor 600, preferably generally shaped as a tube collimator. Radiationquanta 602 are registered by the radiation sensor 600, as describedhereinabove, thereby providing the average number of quanta per unittime. The radiation sensor 600 may be moved around a volume of interest604, preferably, with six-dimensional freedom of motion, and sixdimensional monitoring. However, it will be appreciated that a morelimited motion and corresponding monitoring, for example, only in the Xand ρ directions is possible. The position of the sensor 600 and itsdirection (as well as the position of the investigated volume 604) areassumed to be known at any given moment (FIG. 27A).

The tube collimator is preferably provided with a plane circulardetector 606 of radiation quanta. The quanta detector 606 is preferablydisposed on a rear end 608 of the tube and radiation quanta can reachthe detector 606 only through an open front end 610 of the tube (FIG.27B)

Reference is now made to FIG. 27C, which illustrates a system ofcoordinates (x, y, z) with the origin O in the center of the radiationsensor 600, the (x, y) plane being the plane of the detector, and the zaxis being in the center of the collimator tube. The geometry parametersof the collimator tube—height h and radius ρ—are known.

From the rotational symmetry of the tube, it is clear that having aradiation source Q=Q(x, y, z) with the total intensity I uniformlyradiating in all directions, the portion of the intensity registered bythe quanta detector 606 of the radiation sensor 600 is determined onlyby the distance r from Q to the axis of the collimator (axis z) and thedistance z from Q to the (x, y) plane. In other words, there is afunction (r, z), which is defined only by the collimator parameters ρand h (corresponding expressions from ρ, h, r and z may be easilywritten in explicit form), such that the intensity of the radiation spotQ=Q(x, y, z)=Q(r, z) registered by detector 606 is proportional to (r,z) and to the total intensity I of the radiation spot.

Reference is now made to FIG. 27D. It follows from the foregoingdiscussion, that if instead of one radiation spot there is a radiationdistribution I(Q)=I(Q(r, z)) in a volume V, then the radiationintensity, registered by radiation sensor 600, is proportional (withsome constant not depending on the radiation distribution and the sensorposition) to the following integral:

$\begin{matrix}{\int_{V}^{\;}{{I\left( {Q\left( {r,z} \right)} \right)}{\Phi \left( {r,z} \right)}\ {Q}}} & (1)\end{matrix}$

An algorithm for estimating the intensity distribution I(Q) from thevalues obtained in the measurement scheme of Equation (1) is nowdiscussed. For the sake of simplicity, the first case is discussed withreference to FIG. 27E for a two-dimensional problem, wherein intensitiesI(Q) are distributed in some 2-dimensional plane. The 3D problem is adirect generalization of the corresponding 2D problem, as is explainedhereinbelow.

As seen in FIG. 27E, the radiation sources are distributed in arectangular region V in a plane. Two systems of coordinates areconsidered. The first one is the sensor coordinate system (x, y, z)corresponding to the sensor 600. The second one is the radiation sourcecoordinate system (u, v, w) corresponding to the radiation sources plane(u, v).

It is assumed that at each discrete time increment, the position of theorigin of (x, y, z) system and the direction of the z-axis unit vectorin (u, v, w) coordinates are known. In other words, the position anddirection of the moving sensor in the (u, v, w) coordinate system isknown, and the (u, v, w) coordinate system is assumed to be motionless.

The radiation sources are considered to be distributed in accordancewith the distribution function I(Q)) in some bounded, given rectangle Von the plane (u, v). I(Q)=I(u, v) is the unknown and sought-forradiation (or radiation intensity) distribution function defined in V.

To regularize the problem of estimation of the radiation distributionfunction I(Q), the function I(Q) will be considered to be given fromsome finite dimensional space H of functions defined in V. In otherwords, the function I(Q) itself will not be estimated but rather somefinite dimensional approximation of the distribution I(Q).

The simplest approach to finite dimensional approximation is tosubdivide the rectangle V into sets of equal rectangular cells andconsider the space H of step-functions corresponding to this subdivision(i.e., the space of functions that are constant in the cells ofsubdivision), as shown in FIG. 27F.

If the subdivision of rectangle V into small rectangles is sufficientlyfine, then this step-function approximation is good enough for theestimation of radiation distribution I(Q).

Let each side of rectangle V be divided into n equal parts (FIG. 27F).Then m=n² is the dimension of the space H of step-functions on thecorresponding subdivision.

The space H is naturally isomorphic to the m-dimensional space of n×nmatrixes (with its natural scalar product <•, •>).

Let I=(I_(i j))_(i, j=1, . . . , n) be the unknown element of H that itis desired to estimate. Suppose that element I is measured on Kfunctionals {Φ_(k)}_(k=1 . . . K) of the form of the integral (1):

<I,Φ _(k)>=Σ_(i,j=1 . . . n) I _(ij)Φ_(ij) ^((k))  (2)

where Φ_(k)=(Φ_(i j) ^((k)))_(i, j=1, . . . , n), k=1, . . . , K (afterapproximation of function I(Q), by the corresponding step-function, theintegral (1) is transformed to the sum (2)).

Functionals Φ_(k), k=1, . . . , K, correspond to K discrete positions ofthe sensor (FIG. 27E). Knowing the explicit expressions for functions(r, z) from (1) and knowing for each time moment k, the position of thesensor relative to inspection region V, one can calculate all matrixesΦ_(k)=(Φ_(i j) ^((k)))_(i, j=1, . . . , n), k=1, . . . , K.

Accordingly the following scheme of measurements are obtained:

ψ_(k) =<I,Φ _(k)>+ε_(k) ,k=1, . . . ,K.  (3)

Here ψ_(k) are results of measurements of the unknown element I of thespace H, and ε_(k) are random errors (ε_(k)— independent randomvariables, Eε_(k)=0, k=1, . . . , K).

Let M: H→H the operator in the space H of the form:

M=Σ _(k=1 . . . K)Φ_(k)

Φ_(k).  (4)

Then the best non-biased linear estimate Î of the element I is given bythe formula:

Î=M ⁻¹Ψ,  (5)

where M⁻¹: H→H the inverse operator to the operator M of the form (4),and

Ψ=Σ_(k=1 . . . K)ψ_(k)Ψ_(k),  (6)

(where ψ_(k) are the results of measurements of the form (3)).

One problem of using estimates (5) (besides computational problems ifthe dimension m of the space H is very large) is that the operator M:H→H of the form (3) is “bad invertible”. In other words, the estimationproblem is “ill-posed”. It means that having a noise ε_(k) in themeasurements scheme (3), even if the noise is small, may sometimesresult in a very large estimation error dist(I, Î).

This means that the estimation problem requires additionalregularization. This is a general problem of solving a large set oflinear equations. There are several methods for solving such equations.Below is described one of the known methods for solving such equationsbut numerous other methods are also possible, theses include gradientdecent methods such as in(http://www-visl.technion.ac.il/1999/99-03/www/) and other methods thatare generally known in the art. Further, it is possible to improve theimage reconstruction by taking into account the correlation betweenmeasurements as they are done with substantial overlap. Also, in thefollowing description, a regular step function is assumed for therepresentation of the pixels or voxels, other basis may be used such aswavelet basis, Gaussian basis, etc., which may be better suited for someapplications.

To obtain regularized estimate Î_(R) instead of the estimate Î, theeigenvector decomposition of the operator M may be used:

Let φ₁, φ₂, . . . , φ_(m) be eigenvectors of operator M: H→Hcorresponding to Eigenvalues λ₁≧λ₂≧ . . . ≧λ_(m)≧0.

Let R be some natural number, 1<R<m (R is the “regularizationparameter”). Let H^((R)) be the subspace of the space H spanned by thefirst R eigenvectors (φ₁, . . . φ_(R).

H ^((R)) =sp{φ _(k)}_(k=1 . . . R.)  (7)

Let P^((R)): H→H^((R)) be the orthogonal projection on subspace H^((R)).

The regularized estimate Î_(R) may be obtained as follows:

Let Ψ_(k) ^((R))=P^((R))Ψ_(k), k=1, . . . , K.

Ψ^((R))=Σ_(k=1 . . . K)ψ_(k)Φ_(k) ^((R)),  (8)

M^((R)): H^((R))→H^((R)) the operator of the form:

M ^((R))=Σ_(k=1 . . . K)Φ_(k) ^((R))

Φ_(k) ^((R))  (9)

(operator M^((R)) is the restriction of the operator M of the form (4)to the subspace H^((R)) of the form (7)),

then Î^((R))=(M^((R)))⁻¹ Ψ^((R)).  (10)

When the regularization parameter R is properly chosen (so that theeigenvalue λ_(R) is not too small), then the estimate (10) becomesstable.

There are several possible approaches to choosing the parameter R. Oneapproach is to leave R as a “program parameter” and to obtain thereasonable value “in experiment”. Another approach is to choose some“optimal” value. This is possible if the covariation operators of therandom noise ε_(k) in (3) are known, and information about the element Iof the space H is known a priori.

The subdivision of rectangle V into a large number of equal rectangleshas the disadvantage of making the dimension of the space H too big(especially in the 3D case). If each side of rectangle V is subdividedinto n equal parts, then the dimension of the space H will be n² and thedimension of the matrices used in solving the corresponding estimationequations would be n²×n²=n⁴ (in the 3D case, n³×n³=n⁶). It is clear thatfor large n, this situation may cause serious memory and computationtime.

In accordance with a preferred embodiment of the present invention, anirregular subdivision of the rectangle V is used. This irregularapproach may significantly decrease the dimension of the problem andfacilitate computer calculations.

More specifically, a drawback of the regular subdivision of theinvestigated region V, discussed hereinabove, is that a lot of cellsthat actually have no signal may be taken into account (FIG. 27F). Itwould be much better to have small cells only in regions with highsignal and have big cells in regions with low signal.

Reference is now made to FIG. 27G, which illustrates an advantageousirregular cell subdivision of rectangle V, in accordance with apreferred embodiment of the present invention.

In a first stage, regular subdivisions are made in “large” cells, andmeasurements and estimations are made as described hereinabove. In thismanner, the intensity distribution is estimated in the large cells.

In a second stage, the large cells, which have an intensity larger thansome threshold, are subdivided into 4 equal subcells (or 8 subcells inthe 3D case). A suitable threshold may be obtained by taking the averageintensity (of all large cells) minus two (or three) sigmas (standarddeviation), for example. Measurements and estimations are made in thesesubdivisions as described hereinabove.

The act of subdivision and subsequent measurements and estimations arecontinued until a desired accuracy is reached at some smaller level ofsubdivision, typically defined by the computational and memorycapabilities of the computer being used.

The 3D problem may be treated in the same way as the 2D case, the onlydifference being that instead of rectangle V, there is a parallelepipedV (FIG. 27D). Accordingly, the cells in each subdivision are alsoparallelepipeds.

In accordance with the present invention, the wide-aperturecollimation-deconvolution algorithms described hereinabove may be usedfor a variety of imaging systems. For example, the algorithms may beused with single radiation detector probe, an array of radiationdetector probes, large gamma cameras of various designs, such asmulti-head cameras, general purpose cameras, and automatic white balance(AWB) scanners. The algorithms are suitable for SPECT and planarimaging, and may be used for all types of isotopes of with any type ofphoton energy.

From the foregoing discussion, the skilled artisan will appreciate thatthe algorithms described hereinabove may be used to predict the locationof the radiation source and the uncertainty region (based on the systemmeasurement errors) in the vicinity of the radiation source. Thealgorithms also guide the user to perform additional measurements tominimize the uncertainty region according to the requirements of thesystem operator.

The algorithms thus comprise a feedback system that employs analysis todetermine the bounds of an uncertainty region about the radiationsource, and which guides medical personnel to conduct additional scansin these uncertainty regions to improve accuracy, reduce error, andhence minimize the bounds of the uncertainty regions.

Continuous sampling with radiation probe 202 may provide localization ofa tumor and a physiological radiation activity map of the tumor region.Higher safety and accuracy are gained by a greater number of scans.

Other deconvolution methods are known and are often used in imageprocessing procedures. Examples of such deconvolution methods aredescribed in U.S. Pat. No. 6,166,853 to Sapia et al., the disclosure ofwhich is incorporated herein by reference. (However, it is appreciatedthat these are just examples and the present invention is not limited tothe deconvolution methods mentioned in U.S. Pat. No. 6,166,853)

Reference is now made to FIG. 22 which illustrates a flow chart of aradiation map reconstruction algorithm, by another system ofwide-aperture collimation-deconvolution algorithms, in accordance withanother preferred embodiment of the present invention.

In typical image acquisition, light (or other electromagnetic waveenergy) passes through a finite aperture to an image plane. The acquiredimage is a result of a convolution of the source object's light with theaperture of the imaging system. A system transfer-function may begenerally obtained directly by taking the Fourier transform of theaperture. As is known in the art, the blurring effects due toconvolution generally exist in two-dimensions only, i.e., the x-yplanes. A point-spread-function (PSF) is an expression used to describethe convolutional blurring in two-dimensions. The PSF physically resultsfrom imaging a point source. The Fourier transform of the PSF is thesystem transfer-function, which is obtained by convolving the systemtransfer-function with a Dirac-delta function. A point source is thephysical equivalent of a Dirac-delta function, and, in the frequencydomain, the Dirac-delta function is a unity operator across thespectrum. Therefore, the Fourier transform of the PSF should be theFourier transform of the aperture. However, the PSF contains noise andblurring due to other effects such as aberrations.

The PSF contribution to the overall blurriness may be diminished oreliminated by deconvolution.

In the case of the present invention, the transfer function of theradiation detector may be determined by taking the Fourier transform ofthe aperture of the detector, and taking into account the noise andblurring due to other effects such as aberrations (step 500). An exampleof a transfer function may be a normal distribution. Using knownmathematical techniques, the deconvolution of the transfer function maybe determined (step 502).

The count readings of each spatial location of the detector constitutethe sum of radiation counts from all the voxels (or pixels in the caseof two-dimensional maps, the term “voxel” being used herein to includeboth pixels and voxels) within the detector's field of view. At leastone voxel, or preferably each such voxel, may be assigned a count valuebased on the deconvolution of the unique transfer function of theradiation detector in use (step 504). An additional mathematicalprocedure may treat the various values that each voxel receives due tothe multiple readings from viewpoints of different detectors (step 506).This treatment may constitute for example a simple algebraic average,minimum value or reciprocal of averaged reciprocals in order to producea single value of readings in each voxel. The deconvolution is then usedto reconstruct the voxels of the radiation map with diminished or noblurriness (step 508).

The algorithms described herein are applicable not only to the analysisof readings obtained using a directional radioactivity detector, ratherthey also apply for spatially sensitive (pixellated) radioactivitydetectors. In this case, the readings of each pixel are algorithmicallytreated as described herein like for a directional radioactivitydetector. The motivation behind using a spatially sensitive detector isto save on measurement time by receiving readings from a multitude ofdirections in parallel. This, in essence, creates a number ofoverlapping low resolution images which can then be processed to form ahigh resolution image. In addition, the spatially sensitive detector canbe scanned to improve even further the resolution using the algorithmsdescribed hereinabove.

Thus, the same algorithms that apply for a directional detector applyfor the spatially sensitive detector, only now instead of one radiationreading at each position, a large set of desecrate positions areprocessed in parallel. Each pixel can be seen as a separate detectorwith an angle of acceptance dictated by the geometry of a segmentedcollimator employed thereby. Each of the pixels occupies a differentposition in space and hence can be seen as a new position of a singledirectional probe by the algorithm described herein. It is alsopossible, like with the directional detector, to scan the whole set ofpixels by scanning the spatially sensitive detector and to acquire a newset of data points from the new position. Once obtaining a lowresolution image from each of the pixels of the spatially sensitivedetector, a super resolution algorithm can be employed to generate animage of higher resolution. Suitable super resolution algorithms aredescribed in, for example, J. Acoust. Soc. Am., Vol. 77, No. 2, February1985 Pages 567-572; Yokota and Sato, IEEE Trans. Acoust. Speech SignalProcess. (April 1984); Yokota and Sato, Acoustical Imaging (Plenum,N.Y., 1982, Vol. 12; H. Shekarforoush and R. Chellappa, “Data-DrivenMulti-channel Super-resolution with Application to Video Sequences”,Journal of Optical Society of America-A, vol. 16, no. 3, pp. 481-492,1999; H. Shekarforoush, J. Zerubia and M. Berthod, “Extension of PhaseCorrelation to Sub-pixel Registration”, IEEE Trans. Image Processing, toappear; P. Cheeseman, B. Kanefsky, R. Kruft, J. Stutz, and R. Hanson,“Super-Resolved Surface Reconstruction From Multiple Images,” NASATechnical Report FIA-94-12, December, 1994; A. M. Tekalp, M. K. Ozkan,and M. I. Sezan, “High-Resolution Image Reconstruction forLower-Resolution Image Sequences and Space-Varying Image Restoration,”IEEE International Conference on Acoustics, Speech, and SignalProcessing (San Francisco, Calif.), pp. III-169-172, Mar. 23-26, 1992,http://www-visl.technion.ac.il/1999/99-03/www/, which are incorporatedherein by reference.

Referring further to the drawings, FIGS. 28A-28F schematicallyillustrate a radioactive emission probe 700, preferably, handheld, forfree-hand scanning, in accordance with preferred embodiments of thepresent invention.

As seen in FIG. 28A, handheld probe 700 is a lightweight nuclear imagingcamera, adapted for free-hand movement and having a positioning device714 incorporated thereto.

Handheld probe 700 includes a radiation detector 706, mounted onto ahousing 702, wherein the various electronic components (not shown) arecontained. Housing 702 is preferably formed of a rigid, lightweightplastic, a composite, or the like, and includes a handle 704, for easymaneuvering of probe 700.

A control unit 710, which may be mounted on housing 702, may includebasic control knobs, such as “stop”, “start”, and “pause.” Additionallyor alternatively, control unit 710 may be a computer unit, such as amicrocomputer or the like, having processing and memory units, containedwithin housing unit 702. Additionally, control unit 710 may include adisplay screen 718, mounted on housing unit 702, for displayinginformation such as gamma counts, gamma energies, device position, andthe like. Display screen 718 may be interactive, for control anddisplay. Control unit 710 may further include a data unit 720, forreceiving a diskette, a minidisk, or the like. Preferably, data unit 720is a read and write unit. An eject button 722 may be included with dataunit 720.

A cable 724 may provide signal communication between handheld probe 700and a main-data-collection-and-analysis-system 726, such as a PCcomputer or a server. Additionally or alternatively, signalcommunication with main-data-collection-and-analysis-system 726 may bewireless, for example, by radio frequency or infrared waves.Alternatively, hand-held probe 700 may be a stand-alone unit, operativewith built-in computer unit 710. Additionally, data may be transferredto and from main-data-collection-and-analysis-system 726 via a disketteor a minidisk, from data unit 720.

A preferably rechargeable battery 716 may be attached to housing unit702. Additionally or alternatively, a cable 712 may be used to providepower communication between handheld probe 700 and the grid (not shown),or between handheld probe 700 andmain-data-collection-and-analysis-system 726.

Positioning device 714 is preferably a Navigation sensor, fordetermination of six coordinates X, Y and Z axes and rotational anglesρ, θ and φ, as shown in FIG. 28F. Positioning device 714 may be mountedon housing 702, or located within it. Preferably, positioning device 714is a magnetic tracking and location system, known as miniBIRD,commercially available from Ascension Technology Corporation, P.O. Box527, Burlington, Vt. 05402 USA(http://www.ascension-tech.com/graphic.htm). Alternatively, any otherknown, preferably 6D tracking and positioning system may be used.Positioning device 714 maintains signal communication with either one orboth of control unit 710 and main-data-collection-and-analysis-system726.

As seen in FIGS. 28A and 28B, nuclear imaging detector 706 may be adisk, forming a single pixel, having a diameter D of about 10 mm, and alength L1 of about 5 mm. It will be appreciated that other dimensions,which may be larger or smaller, are also possible. Preferably, radiationdetector 706 is connected to a preamplifier (not shown).

Additionally, as seen in FIGS. 28A and 28B, handheld probe 700 mayinclude a tubular, wide-bore collimator 708, formed for example, of leador of tungsten. The collimator may have a length L2 of about 10 to 30mm, and preferably, about 15 mm. It will be appreciated that otherdimensions, which may be larger or smaller, are also possible.

Alternatively, as seen in FIGS. 28C and 28D, collimator 708 may be asquare grid of 4×4 cells 705, or 16×16 cells 705. Collimator grid 708may have sides w of about 10 mm, and length L2 of about 20-40 mm mm. Itwill be appreciated that other dimensions, which may be larger orsmaller, are also possible.

As seen in FIG. 28E, each cell 705 may include a plurality of radiationdetector pixels 703, formed for example, as square pixels of 5×5 mm andlength L1 also of about 5 mm. The reason for the division of nucleardetector 708 into pixels 703 is that while the wide aperture of a coarsecollimator increases the counting efficiency, radiation detectoruniformity is better maintained when the size of the radiation detectorsis small. Thus, it is desired to combine a coarse collimator grid with afine pixel size. Lead spacers are placed between the pixels. Fromspatial resolution consideration, each cell 705 preferably operates as asingle pixel, but each pixel 703 within cell 705, must be calibratedindividually, to correct for material nonuniformity, as will bedescribed hereinbelow, in conjunction with FIG. 29B.

Preferably, radiation detector 706 is equipped with a positive contactarray, for forming contact with each pixel 703, and a common negativecontact. Preferably, each pixel 703 is connected to a preamplifier.

Preferably, radiation detector 706 is a single module array, forexample, of 4×4, or of 16×16 pixels, of room temperature CdZnTe,obtained, for example, from IMARAD IMAGING SYSTEMS LTD., of Rehovot,ISRAEL, 76124, www.imarad.com.

Room temperature solid-state CdZnTe (CZT) is among the more promisingnuclear detectors currently available. It has a better count-ratecapability than other detectors on the market, and its pixilatedstructure provides intrinsic spatial resolution. Furthermore, because ofthe direct conversion of the gamma photon to charge-carriers, energyresolution is enhanced and there is better rejection of scatter eventsand improved contrast.

In accordance with the present invention, the detector may be optimized,in accordance with the teaching of “Electron lifetime determination insemiconductor gamma detector arrays,” http://urila.tripod.com/hecht.htm,“GdTe and CdZnTe Crystal Growth and Production of Gamma RadiationDetectors,” http://members.tripod.com/-urila/crystal.htm, and “DrivingEnergy Resolution to the Noise Limit in Semiconductor Gamma DetectorArrays,” Poster presented at NSS2000 Conference, Lyon France, 15-20 Oct.2000, http://urila.tripod.com/NSS.htm, all by Uri Lachish, of GumaScience, P.O. Box 2104, Rehovot 76120, Israel, urila@internet-zahav.net,all of whose disclosures are incorporated herein by reference.

Accordingly, radiation detector 706 may be a monolithic CdZnTe crystal,doped with a trivalent donor, such as indium. Alternatively, aluminummay be used as the trivalent donor. When a trivalent dopant, such asindium, replaces a bivalent cadmium atom within the crystal lattice, theextra electron falls into a deep trap, leaving behind an ionized shallowdonor. The addition of more donors shifts the Fermi level from below thetrapping band to somewhere within it. An optimal donor concentration isachieved when nearly all the deep traps become occupied and the Fermilevel shifts to just above the deep trapping band.

Optimal spectral resolution may be achieved by adjusting the gammacharge collection time (i.e., the shape time) with respect to theelectron transition time from contact to contact. Gamma photons areabsorbed at different depth within the detector where they generate theelectrons. As a result, these electrons travel a different distance tothe counter electrode and therefore produce a different external signalfor each gamma absorption event. By making the shape time shorter thanthe electron transition time, from contact to contact, these externalsignals become more or less equal leading to a dramatic improvement inresolution.

Furthermore, for a multi-pixel detector, the electrons move from thepoint of photon absorption towards the positive contact of a specificpixel. The holes, which are far slower, move towards the negativecontact, and their signal contribution is distributed over a number ofpixels. By adjusting the gamma charge collection time (i.e., the shapetime) with respect to the electron transition time from contact tocontact, the detector circuit collects only the electrons' contributionto the signal, and the spectral response is not deteriorated by thecharge of the holes.

For an optimal detector, crystal electrical resistively may be, forexample, about 5×10 ⁸ ohm cm. The bias voltage may be, for example, −200volts. The shape time may be, for example, 0.5×10⁻⁶ sec. It will beappreciated that other values, which may be larger or smaller are alsopossible.

In accordance with a preferred embodiment of the present invention,control unit 710 includes data acquisition and control components, whichfurther include a very small, single-module Detector Carrier Board(DCB). The DCB preferably includes a temperature control system, havinga temperature sensor, a desired temperature setting, and a heat removalsystem, preferably operative by peltier cooling. Preferably, the DCBfurther includes a High Voltage (HV) power supply, a HV setting, a HVand dark current output-sensing, and power regulators, preferably havinga range of about ±2V, a XAIM_MBAIS current driver, and noise filteringcomponents.

Preferably, data acquisition and control components are based onApplication Specific Integrated Circuit (ASIC) architecture for thecontrol, calibration, and readout. ASIC data acquisition and controlincludes calibration of ASIC data, I/V conversion, amplification andshaping. Data identification includes threshold discrimination (i.e.,noise rejection) and energy windows discrimination, wherein there may beone or several energy windows per radionuclide. Preferably, there are upto four energy windows per radionuclide, but it will be appreciated thatanother number is also possible.

Data acquisition relates to counting events per pixel, per energywindow, for an acquisition period, which may be very short, for example,about 100 msec, in a continuous mode. It will be appreciated that largeror smaller counting periods are also possible.

The events per pixel 703, per energy window, is synchronized withsignals from positioning device 714, preferably via a DIO card, eitherwithin control unit 710, or main-data collection-and-analysis-system726, or both. Additionally, initialization information, includingDigital to Analog Convertor (DAC) information, may be stored in softwareof main-data-collection-and-analysis-system 726 or within control system710.

Preferably, an Integrated Drive Electronics (IDE) system provides theelectronic interface.

Referring further to the drawings, FIGS. 29A-29B schematicallyillustrate the manner of calibrating handheld probe 700, in accordancewith preferred embodiments of the present invention. Non-uniformity ofthe material leads to different sensitivities for each pixel. Acorrection is required to eliminate the sensitivity effect of thedifferent pixels.

An energy spectrum for each pixel 703 of radiation detector 706 isobtained in a classical fashion. Radiation detector 706 is irradiated bya flood source of a known photon energy, for example, 24 KeV, and foreach pixel 703, such as 703(1), 703(2), 703(3). 703(n), a summation ofcounts as a function of voltage is obtained, during a finite timeinterval. The location of the energy peak, corresponding to the peakphoton energy, for example, 24 KeV is bounded, by an upper level (UL)and a lower level (LL) for that pixel 703(n), forming an energy windowfor the specific pixel. Only photons which fall within the boundedenergy window for each pixel are counted, to exclude Compton scatteringand pair production. A multi-energy source may be used, and severalenergy windows may be bounded for each pixel.

After the energy windows, representing specific energy peaks, aredetermined for each pixel, they are stored in a memory unit.

During acquisition, each collimator cell 705(n) is operative as a singlepixel. Thus, the sensitivity correction factor has to be at acollimator-cell level, rather than at a pixel-level.

To calculate the sensitivity correction factor, an average counts percell 705, is obtained as:

${{average}\mspace{14mu} {counts}\mspace{14mu} {per}\mspace{14mu} {cell}\mspace{14mu} 705} = \frac{{sum}\mspace{14mu} {of}\mspace{14mu} {counts}\mspace{14mu} {in}\mspace{14mu} {all}\mspace{14mu} {pixels}\mspace{14mu} 703\mspace{14mu} {of}\mspace{14mu} {detector}\mspace{14mu} 706}{{number}\mspace{14mu} {of}\mspace{14mu} {collimator}\mspace{14mu} {cells}\mspace{14mu} 705\mspace{14mu} {in}\mspace{14mu} {detector}\mspace{14mu} 706}$${{correction}\mspace{14mu} {factor}\mspace{14mu} {for}\mspace{14mu} {cell}\mspace{14mu} 705\; (n)} = \frac{{average}\mspace{14mu} {counts}\mspace{14mu} {per}\mspace{14mu} {cell}{\mspace{14mu} \;}705}{{counts}\mspace{14mu} {in}\mspace{14mu} a\mspace{14mu} {cell}\mspace{14mu} 705\; (n)}$

During acquisition, the sum of counts within the bounded windows of allpixels in cell 705(n) is multiplied by the correction factor for thatcell 705(n) to obtain a sensitivity corrected count rate.

Referring further to the drawings, FIG. 30 schematically illustrates themanner of synchronizing event readings by radiation detector 706 andposition readings by positioning device 714, in accordance withpreferred embodiments of the present invention. Preferably, eventreadings, for each cell 705 is tabulated with the position reading, foreach time interval, such as t₁, t₂, t₃, after correction forsensitivity.

Referring further to the drawings, FIGS. 31A-31C describe a spatialresolution bar-phantom test of probe 700, in accordance with a preferredembodiment of the present invention. As in FIGS. 17 and 18, hereinabove,the aim of the bar phantom test is to evaluate image quality bymeasuring image resolution of a bar phantom. Spatial resolution isexpressed in terms of the finest bar pattern, visible on nuclear images.

As seen in FIG. 31A, a bar phantom 750 was constructed of lead strips752, encased in a plastic holder (not shown). The lead strips were ofequal width and spacing. Thus, a 2-mm bar pattern consisted of 2-mm widestrips 752, separated by 2-mm spaces 754, along an x axis. The thicknessof lead strips 752 was, sufficient to ensure that bars 752 are virtuallyopaque to the γ rays being imaged.

A uniform radiation field, formed of a sheet source (not shown), wasplaced directly behind the bar phantom 750, and bar phantom 750 wasscanned with probe 700.

Probe 700 of the present embodiment comprised an eV Pen-type CdZnTesolid state detector with collimator diameter D, of 5 mm, and collimatorlength, L2 of 30 mm. The CdZnTe detector was attached to a chargesensitive preamp probe and connected to a Digital and analog V-targetalpha system PC cards. The probe energy window calibration was performedat the beginning of the test by using EG&G Ortec Maestro 2K Tramp ISASpectra Processing connected to the analog and the digital V-targetalpha system cards. The probe position information was provided byMicroScribe-6DLX arm manufactured by Immersion Corporation. Acquisitiontime was 6 minutes.

FIG. 31B illustrates a bar phantom image 760, produced by probe 700,after 6 minutes of acquisition time, in accordance with the presentinvention. Image 760 shows bar images 756, and space images 758, of the2-mm bar pattern, clearly demonstrating a high degree of spatialresolution.

FIG. 31C provides a similar observation, based counts as a function oflength axis, X. The well-defined, narrow sinusoidal peaks of counts as afunction of length axis demonstrate a high degree of spatial resolutionfor the 2-mm bar pattern.

Referring further to the drawings, FIGS. 32A-32D describe a spatialresolution bar-phantom test of a prior-art probe, a High ResolutionGamma Camera of a competitor. Although for the prior art probe,acquisition time was 3 hours, a bar pattern is barely observed for a3-mm bar pattern of FIG. 32C, and clearly observed only for a 3.5 mm barpattern of FIG. 32D.

Table 1 provides a comparison between the prior art technology and thatof the present invention, illustrating a clear advantage to the presentinvention.

TABLE 1 Detector - Maximum Imaging collimator Acquisition SpatialTechnology arrangement Time Resolution a multi-pixel 180 min. Worse thenHigh-resolution detector, 3 mm Gamma-Camera tube collimators of a 1.9 mmin dia. competitor's 40 mm in length Probe 700 single-pixel  6 min. 2 mmof FIG. 28A detector 10 mm in dia. tube collimator 25 mm in lengthSummary V-target V-target's probe resolution is 30 time is 2 time fasterbetter

Good energy resolution of the detector is important for preciseidentification and separation of γ rays, for radionuclide identificationand for scatter rejection. A major advantage of CdZnTe detector isbetter energy resolution, for example than that of the Nai(Tl) detector.

Referring further to the drawings, FIGS. 33A-33B illustrate the energyresolution of a single pixel 703 of probe 700, in accordance with thepresent invention.

FIG. 33A compares a spectrum of a single pixel of a CdZnTe detector withthat obtained by a single pixel of an NaI(Tl) detector, at ˜120 KeV,showing the difference in energy resolution between the two detectors.Clearly, the a CdZnTe detector provides a sharper, more well-definedenergy peak, which is about ⅓ higher than the Nai(Tl) peak, and having aFWHM that is ⅔ that of the Nai(Tl) peak.

FIG. 33B provides a spectrum of a a CdZnTe detector, showing good energyresolution at the low energy level of about 24 KeV.

Referring further to the drawings, FIGS. 34A-34C schematicallyillustrate endoscopic radioactive emission probes 800, adapted to beinserted into the body, on a shaft or a catheter 713, via a trocar valve802 in a tissue 810.

As seen in FIG. 34A, endoscopic radioactive emission probe 800 mayinclude, for example, radiation detector 706, for example, of a singlepixel, wide-bore collimator 708, position tracking system 714, and anextracorporeal control unit 804. Preferably, the diameter of radiationdetector 706, D_(detector), is less than 12 mm. In accordance withanother embodiment of the present invention, radioactive emission probe800 may be used in body lumens, for example, for scanning the stomach orthe uterus.

As seen in FIG. 34B, endoscopic radioactive emission probe 800 mayinclude several radiation detectors 706, for example, of a single pixeleach, and having a dedicated collimator 708, each. Unlike the collimatorgrid structure of FIGS. 28C and 28D, the collimators of FIG. 34B are notparallel to each other, and each may point in a different direction.Alternatively, each radiation detector 706 may include a plurality ofpixels.

As seen in FIG. 34C, endoscopic radioactive emission probe 800 mayinclude several radiation detectors 706, which may include wide-anglecollimators.

Referring further to the drawings, FIG. 35 schematically illustrates amanner of calculating a distance, d, between radioactive emission probe,such as probe 700 (FIGS. 28A-28F) and a radiation source 766, based onthe attenuation of photons of different energies, which are emitted fromthe same source. Probe 700 is on an extracorporeal side 762 andradiation source 766 is on an intracorporeal side 764 of a tissue 768.

The transmitted intensity of a monochromatic gamma-ray beam of photonenergy E and intensity I₀ crossing an absorbing layer, such as a tissuelayer, of a depth d, is a function of the photons energy E, the layerthickness d, and the total linear attenuation coefficient as a functionof energy, μ_(total), of the material constituting the layer. Thetransmitted intensity is given by:

I(E)=I ₀(E)exp(−μ_(total)(E)d)  (Eq-1)

The linear attenuation coefficient is expressed in cm⁻¹ and is obtainedby multiplying the cross section (in cm²/gr) with the absorbing mediumdensity. The total linear attenuation coefficient accounts forabsorption due to all the relevant interaction mechanisms: photoelectriceffect, Compton scattering, and pair production.

The layer thickness d, which is the depth of the radiation source, canbe measured using a radionuclide that emits two or more photon energies,since attenuation is a function of the photon energy.

For photons of different energies, originating from a same source, andhaving the same half-life, we get:

$\begin{matrix}{\frac{I\left( E_{1} \right)}{I\left( E_{2} \right)} = {\frac{I_{0}\left( E_{1} \right)}{I_{0}\left( E_{2} \right)} \cdot \frac{^{{- {\mu {(E_{1})}}} \cdot d}}{^{{- {\mu {(E_{2})}}} \cdot d}}}} & \left( {{Eq}\text{-}2} \right) \\{A = {{{\frac{I\left( E_{1} \right)}{I\left( E_{2} \right)}\&}B} = \frac{I_{0}\left( E_{1} \right)}{I_{0}\left( E_{2} \right)}}} & \left( {{Eq}\text{-}3} \right) \\{R = {\frac{A}{B} = \frac{\left\lbrack {{I\left( E_{1} \right)}/{I\left( E_{2} \right)}} \right\rbrack}{\left\lbrack {{I_{0}\left( E_{1} \right)}/{I_{0}\left( E_{2} \right)}} \right\rbrack}}} & \left( {{Eq}\text{-}4} \right) \\{A = {\left. {B \cdot \frac{^{{- {\mu {(E_{1})}}} \cdot d}}{^{{- {\mu {(E_{2})}}} \cdot d}}}\Rightarrow R \right. = \frac{^{{- {\mu {(E_{1})}}} \cdot d}}{^{{- {\mu {(E_{2})}}} \cdot d}}}} & \left( {{Eq}\text{-}5} \right) \\{R = {\left. ^{{\{{{\mu {(E_{2})}} - {\mu {(E_{1})}}}\}}d}\Rightarrow{\ln (R)} \right. = {\left\lbrack {{\mu \left( E_{2} \right)} - {\mu \left( E_{1} \right)}} \right\rbrack d}}} & \left( {{Eq}\text{-}6} \right)\end{matrix}$

The source depth, d, can be calculated as:

$\begin{matrix}{d = {\frac{\ln (R)}{{\mu \left( E_{2} \right)} - {\mu \left( E_{1} \right)}} = \frac{\ln \left\{ {\left\lbrack {{I\left( E_{1} \right)}/{I\left( E_{2} \right)}} \right\rbrack/\left\lbrack {{I_{0}\left( E_{1} \right)}/{I_{0}\left( E_{2} \right)}} \right\rbrack} \right\}}{{\mu \left( E_{2} \right)} - {\mu \left( E_{1} \right)}}}} & \left( {{Eq}\text{-}7} \right)\end{matrix}$

The linear attenuation coefficients for specific energies are calculatedfrom the photon mass attenuation and energy absorption coefficientstable.

The ratio I₀(E₁)/I₀(E₂) for a known isotope or isotopes is calculatedfrom an isotope table and the ratio I(E₁)/I(E₂) is measured at theinspected object edges.

Additionally, X-ray information, for example, mammography may be used,for more material linear attenuation coefficient, since X-ray imagingprovides information about the material density.

For example, we may consider a lesion inside a soft tissue, such as abreast, at a distance d that holds an isotope of ¹²³I, which emits twomain photons, as follows:

E₁=27 keV with 86.5%; and

E₂=159 keV with 83.4%.

The soft tissue density is 0.8 g/cm³.

The linear attenuation coefficients for E₁ and E₂ are calculated from J.H Hubbell, “Photon Mass Attenuation and Energy-absorption Coefficientsfrom 1 keV to 20 MeV”, Int. J. Appl. Radiat. Isat. Vol. 33. pp. 1269 to1290, 1982, as follows:

μ(E₁)=0.474 cm/g→μ(E₁)=0.474 cm/g·0.8 g/cm³=0.38 1/cm

μ(E₂)=0.1489 cm/g→μ(E₂)=0.1489 cm/g·0.8 g/cm³=0.119 1/cm

The measured ratio between I(E₁) to I(E₂) on the attenuating soft tissueedge is 0.6153.

Using Eq-7, one gets that the lesion depth is: d=2 cm.

Alternatively, if the measured ration between I(E₁) to I(E₂) is 0.473the extracted lesion depth is: d=3 cm.

It will be appreciated that when two or more isotopes used for obtainingthe different photons, corrections for their respective half-lives maybe required.

It will be appreciated that by calculating the depth of a radiationsource at each position, one may obtain a radiation source depth map, ineffect, a three-dimensional image of the radiation source. Thisinformation may be superimposed on the radiation source image, asproduced by other methods of the present invention, as an independentcheck of the other methods. For example, a radiation source imageproduced by the wide-aperture collimation-deconvolution algorithms maybe superimposed on a radiation source image produced by depthcalculations of based on the attenuation of photons of differentenergies.

EXPERIMENTAL RESULTS

In a series of clinical experiments, some of the basic concepts of theinvention have been tested on patients who were pre-injected with asuitable radiopharmaceutical for their particular pathology.Two-dimensional color-coded maps have been constructed based on a scanof a pre-determined lesion area by a hand-held radiation detector with amagnetic position-tracking system. The resulting maps, which representedthe radiation count level, were compared to images of conventional gammacamera. The list of radiopharmaceuticals tested includes ¹⁸FDG,^(99M)Tc-MDP, ^(99M)Tc sodium pertechnetate, ^(99M)Tc erythrocytes.Similar radiolabeled patterns were observed in the images produced bythe system of the invention and in the images produced by a conventionalgamma camera in the following pathologies:

FIGS. 23A and 23B illustrate radiolabeled patterns observed in imagesproduced by the system of the invention and by a conventional gammacamera, respectively, of an autonomous adenoma of a thyroid of a 58year-old male.

FIGS. 24A and 24B illustrate radiolabeled patterns observed in imagesproduced by the system of the invention and by a conventional gammacamera, respectively, of suspected Paget's disease of a humerus in an 89year-old female.

FIGS. 25A and 25B illustrate radiolabeled patterns observed in imagesproduced by the system of the invention and by a conventional gammacamera, respectively, of chronic osteomyelitis in a 19 year-old female.

FIGS. 26A and 26B illustrate radiolabeled patterns observed in imagesproduced by the system of the invention and by a conventional gammacamera, respectively, of skeletal metastasis from medulloblastoma in an18 year-old male.

FIG. 36 illustrates a control volume 701, and a three-dimensional image703 of the radioactivity emitting source, produced by free-hand scanningof a cancerous prostate gland, ex vivo, with hand-held probe 700 of FIG.28A. The diameter of the detector was 10 mm, and the length of thecollimator was 25 mm. The wide-aperture collimation-deconvolutingalgorithm was used, for image reconstruction. A scale 705 illustratesbrightness as a function of count rate density. The brighter the area,the higher the count density. The Figure shows the capability ofhand-held probe 700 to detect cancerous tissue. The By comparison, thecompetitor's High-resolution, Gamma-Camera failed to detect anycancerous growth.

The following provides a list of known procedures which can takeadvantage of the system and method of the present invention:

In cancer diagnosis the system and method of the present invention canfind uses for screening for cancer and (or) directing invasive diagnosis(biopsies) either from outside the body or by way of endoscopicapproach. Examples include, but are not limited to, lung cancer biopsy,breast cancer biopsy, prostate cancer biopsy, cervical cancer biopsy,liver cancer biopsy, lymph node cancer biopsy, thyroid cancer biopsy,brain cancer biopsy, bone cancer biopsy, colon cancer biopsy, gastrointestine cancer endoscopy and biopsy, endoscopic screening for vaginalcancer, endoscopic screening for prostate cancer (by way of the rectum),endoscopic screening for ovarian cancer. (by way of the vagina),endoscopic screening for cervical cancer (by way of the vagina),endoscopic screening for bladder cancer (by way of the urinary track),endoscopic screening for bile cancer (by way of the gastrointestinaltrack), screening for lung cancer, screening for breast cancer,screening for melanoma, screening for brain cancer, screening for lymphcancer, screening for kidney cancer, screening for gastro intestinalcancer (from the outside).

In the special case of MRI, the radiation detector can be combined andpackaged together with a small RF coil for the transmission andreception or reception only of the MRI signals in a rectal probeconfiguration for prostate diagnosis and treatment or any other closeconfinement position such as the vagina, airways, the uper portion ofthe gastrointestinal track, etc)

Procedures known as directing localized treatment of cancer can alsobenefit from the system and method of the present invention. Examplesinclude, but are not limited to, intra tumoral chemotherapy, intratumoral brachytherapy, intra tumoral cryogenic ablation, intra tumoralradio frequency ablation, intra tumoral ultrasound ablation, and intratumoral laser ablation, in cases of, for example, lung cancer, breastcancer, prostate cancer, cervical cancer, liver cancer, lymph cancer,thyroid cancer, brain cancer, bone cancer, colon cancer (by way ofendoscopy through the rectum), gastric cancer (by way of endoscopythrough the thorax), thoracic cancer, small intestine cancer (by way ofendoscopy through the rectum or, by way of endoscopy through thethorax), bladder cancer, kidney cancer, vaginal cancer and ovariancancer.

In interventional cardiology the following procedures can take advantageof the present invention wherein the method and system can be used toassess tissue perfusion, tissue viability and blood flow intraoperatively during PTCA procedure (balloon alone or in conjunction withthe placement of a stent), in cases of cardiogenic shock to asses damageto the heart, following myocardial infarct to asses damage to the heart,in assessing heart failure condition tissue in terms of tissue viabilityand tissue perfusion, in intra vascular tissue viability and perfusionassessment prior to CABG operation.

The radioactivity detector can be mounted on a catheter that is enteredthrough the blood vessels to the heart to evaluate ischemia from withinthe heart in order to guide ablation probes or another type of treatmentto the appropriate location within the heart. Another application whichmay benefit from the present invention is the localization of bloodclots. For example, a radioactivity detector as described herein can beused to asses and differentiate between new clots and old clots. Thus,for example, the radioactivity detector can be placed on a very smallcaliber wire such as a guide wire that is used during PTCA in order toimage blood-vessel clots. Blood-vessel clots can be searched for in theaortic arc as clots therein are responsible for about 75% of strokecases.

Using the method and system of the present invention to assess tissueperfusion, tissue viability and blood flow intra operatively can also beemployed in the following: during CABG operation to asses tissueviability, to mark infarct areas, during CABG operations to asses thesuccess of the re vascularization.

The present invention has many other applications in the direction oftherapeutics, such as, but not limited to, implanting brachytherapyseeds, ultrasound microwave radio-frequency cryotherapy and localizedradiation ablations.

It will be appreciated that many other procedures may also takeadvantage of the present invention.

It is appreciated that certain features of the invention, which are, forclarity, described in the context of separate embodiments, may also beprovided in combination in a single embodiment. Conversely, variousfeatures of the invention, which are, for brevity, described in thecontext of a single embodiment, may also be provided separately or inany suitable subcombination.

Although the invention has been described in conjunction with specificembodiments thereof, it is evident that many alternatives, modificationsand variations will be apparent to those skilled in the art.Accordingly, it is intended to embrace all such alternatives,modifications and variations that fall within the spirit and broad scopeof the appended claims. All publications in printed or electronic form,patents and patent applications mentioned in this specification areherein incorporated in their entirety by reference into thespecification, to the same extent as if each individual publication,patent or patent application was specifically and individually indicatedto be incorporated herein by reference. In addition, citation oridentification of any reference in this application shall not beconstrued as an admission that such reference is available as prior artto the present invention.

What is claimed is:
 1. An intracorporeal-imaging head, comprising: ahousing, which comprises: at least one radioactive-emission probe,mounted on said housing, adapted to image radioactive-emission from atleast two different viewing angles of a same portion of a tissue withoutmovement of the housing.
 2. The intracorporeal imaging head of claim 1,further comprising an imaging system adapted to image said portion. 3.The intracorporeal imaging head of claim 2, wherein the imaging systemis one of a fluoroscope, a computed tomographer, a magnetic resonanceimager, an ultrasound imager, an impedance imager, and an opticalcamera.
 4. The intracorporeal imaging head of claim 1 wherein the atleast one radioactive-emission probe is at least one wide-anglecollimator probe.
 5. The intracorporeal imaging head of claim 4 whereinthe wide-angle collimator probe has a viewing angle of between 81° and280°.
 6. The intracorporeal imaging head of claim 4, further comprising:a data processor for correcting blurring effects from said wide-anglecollimator probe.
 7. The intracorporeal imaging head of claim 6, whereinsaid wide-angle collimator probe is adapted to obtain image data at afirst resolution or less, and wherein said data processor is furtheradapted for forming a three-dimensional model of said portion of saidtissue, wherein said three-dimensional model is in a second resolution,higher than said first resolution.
 8. The intracorporeal imaging head ofclaim 6, wherein said data processor corrects the blurring effects usinga collimation-deconvolution algorithm based on account estimation of atransfer function.
 9. The intracorporeal imaging head of claim 6,wherein said data processor corrects the blurring effects by calculatinga value of readings for each voxel based on multiple readings from viewpoints of said probe; and using a collimation-deconvolution algorithm toreconstruct voxels of the radiation map with diminished or noblurriness.
 10. The intracorporeal-imaging head of claim 1, furthercomprising: a position tracking system in a fixed positional relationwith said at least one radioactive-emission probe, for providingpositional information for said at least one radioactive-emission probe.11. The intracorporeal imaging head of claim 10 wherein the positiontracking system is one of a fluoroscope, a computed tomographer, amagnetic resonance imager, an ultrasound imager, an impedance imager,and an optical camera.
 12. The intracorporeal imaging head of claim 10,wherein the position tracking system comprises an imaging system. 13.The intracorporeal imaging head of claim 10 wherein the positiontracking system is adapted to track acoustic electromagnetic radiationor magnetic fields.
 14. The intracorporeal imaging head of claim 10wherein the at least one radioactive-emission probe is at least onewide-angle collimator probe.
 15. The intracorporeal imaging head ofclaim 10 wherein the wide-angle collimator probe has a viewing angle ofbetween 81° and 280°.
 16. The intracorporeal imaging head of claim 10,wherein said wide-angle collimator probe is adapted to obtain image dataat a first resolution or less, and wherein the intracorporeal imaginghead further comprises a data processor for forming a three-dimensionalmodel of said portion of said tissue by data processing said image dataat said first resolution and positional information received from saidposition tracking system, wherein said three-dimensional model is in asecond resolution, higher than said first resolution.
 17. Theintracorporeal imaging head of claim 1, wherein said radioactive probecomprises a plurality of radiation detectors.
 18. The intracorporealimaging head of claim 1, wherein said radioactive probe comprises aplurality of single pixel detector units configured to generate imagedata from each of said pixels.
 19. A flexible probe, comprising: anintracorporeal imaging head according to claim 1; and an ultrasonicimager sized for rectal insertion for imaging of prostate.
 20. Aflexible probe, comprising: an intracorporeal imaging head according toclaim 1; and an optical imager sized for rectal insertion for imaging ofcolon.
 21. A device for obtaining an image of a radioactivity emittingsource in a system-of-coordinates, the device comprising: a moveableradioactive emission detector; said radioactive emission detector beingmoveable to perform a plurality of radioactivity measurements of aradioactivity emitting source from a plurality of locations anddirections; a position tracking device which monitors a position of saidradioactive emission detector in relation to said system of coordinates;and a data processor which reconstructs an image of the radioactivityemitting source from a plurality of said radioactivity measurements witha varying spatial resolution and obtains a positional distribution ofsaid radioactive emitting source in relation to said system ofcoordinates using measurements of a same area from multiple locationshaving different resolution according to said position.
 22. The deviceof claim 21, wherein said position tracking device is comprised of aplurality of accelerometers, or a plurality of potentiometers, or issound wave based or radio frequency based, or electromagnetic fieldbased or optically based.
 23. The device of claim 21, wherein saidposition tracking device is responsive to a change in the position ofsaid detector.
 24. The device of claim 21, wherein said detector iscarried by a moveable arm, and said position tracking device isresponsive to movement of said arm.
 25. The device of claim 24, whereinsaid moveable arm is articulated.
 26. The device of claim 21, whereinsaid detector performs said plurality of radioactivity measurements atdifferent resolutions.
 27. The device of claim 21, wherein said dataprocessor reconstructs said image according to radioactivitymeasurements having different resolutions.
 28. The device of claim 27,wherein said resolutions varies according to the distance of saiddetector from said radiation source.
 29. The device of claim 28, whereinsaid spatial resolution is higher for radioactivity measurements as thedetector approaches the radiation source.
 30. The device of claim 21,wherein said radioactive emission detector is configured for free-handscanning, and the position of the detector is monitored by said positiontracking device in a first system-of-coordinates, and furthercomprising: a surgical instrument whose position is monitored by anadditional position tracking device in a second system-of-coordinates;and at least one data processor designed and configured to receive datainputs from said position tracking device, from said radioactiveemission detector and from said additional position tracking device andto calculate the position of the surgical instrument and theradiopharmaceutical up-taking portion of the body component in a commonsystem-of-coordinates.
 31. The device of claim 30, wherein said firstposition tracking device and said additional position tracking deviceare comprised in a single unit.
 32. The device of claim 30, furthercomprising a display device which serves for visual co-presentation ofthe position of said surgical instrument and the radiopharmaceuticalup-taking portion of the body component.
 33. The device of claim 30,wherein said surgical instrument is selected from the group consistingof laser probe, a cardiac catheter, an angioplastic catheter, anendoscopic probe, a biopsy needle, an ultrasonic probe, fiber opticscopes, aspiration tubes, a laparoscopy probe, a thermal probe and asuction/irrigation probe.
 34. The device of claim 30, wherein saidcommon system of coordinates is a three dimensional space.
 35. Thedevice of claim 1, wherein said system of coordinates is a threedimensional space.
 36. The device of claim 1, wherein said dataprocessor is responsive to said plurality of said radioactivitymeasurements to reconstruct a three dimensional image of saidradioactivity emitting source.